Acoustic perfusion devices

ABSTRACT

Methods are disclosed for separating beads and cells from a host fluid. The method includes flowing a mixture containing the host fluid, the beads, and the cells through an acoustophoretic device having an ultrasonic transducer including a piezoelectric material driven by a drive signal to create a multi-dimensional acoustic standing wave. A drive signal is sent to drive the at least one ultrasonic transducer to create the multi-dimensional acoustic standing wave. A recirculating fluid stream having a tangential flow path is located substantially tangential to the standing wave and separated therefrom by an interface region. A portion of the cells pass through the standing wave, and the beads are held back from the standing wave in the recirculating fluid stream at the interface region. Also disclosed is an acoustophoretic device having a coolant inlet adapted to permit the ingress of a cooling fluid into the device for cooling the transducer.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent applicationSer. No. 15/420,073, filed Jan. 30, 2017, which claims priority to U.S.Provisional Patent Application Ser. No. 62/288,500, filed Jan. 29, 2016,and to U.S. Provisional Patent Application Ser. No. 62/296,685, filedFeb. 18, 2016, and to U.S. Provisional Patent Application Ser. No.62/307,934, filed Mar. 14, 2016. U.S. patent application Ser. No.15/420,073 is also a continuation-in-part of U.S. patent applicationSer. No. 15/139,187, filed Apr. 26, 2016, which is acontinuation-in-part of U.S. patent application Ser. No. 14/975,307,filed Dec. 18, 2015, which claims priority to U.S. Provisional PatentApplication Ser. No. 62/256,952, filed on Nov. 18, 2015, and to U.S.Provisional Patent Application Ser. No. 62/243,211, filed on Oct. 19,2015, and to U.S. Provisional Patent Application Ser. No. 62/211,057,filed on Aug. 28, 2015, and to U.S. Provisional Patent Application Ser.No. 62/093,491, filed on Dec. 18, 2014. U.S. patent application Ser. No.14/975,307 is also a continuation-in-part of U.S. patent applicationSer. No. 14/175,766, filed on Feb. 7, 2014, which claims priority toU.S. Provisional Patent Application Ser. No. 61/761,717, filed on Feb.7, 2013, and is also a continuation-in-part of U.S. patent applicationSer. No. 14/026,413, filed on Sep. 13, 2013, which claims the benefit ofU.S. Provisional Patent Application Ser. No. 61/708,641, filed on Oct.2, 2012. U.S. patent application Ser. No. 14/026,413 is also acontinuation-in-part of U.S. Ser. No. 13/844,754, filed Mar. 15, 2013,which claims the benefit of U.S. Provisional Patent Application Ser. No.61/611,159, filed Mar. 15, 2012, and of U.S. Provisional PatentApplication Ser. No. 61/611,440, filed Mar. 15, 2012, and of U.S.Provisional Patent Application Ser. No. 61/754,792, filed Jan. 21, 2013,and of U.S. Provisional Patent Application Ser. No. 61/708,641, filed onOct. 2, 2012. These applications are incorporated herein by reference intheir entireties.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

This invention was made with government support under Grant No.IIP-1330287 (Amendment 003, Proposal No. 1458190) awarded by theNational Science Foundation. The government has certain rights in theinvention.

BACKGROUND

The field of biotechnology has grown tremendously in the last 20 years.This growth has been due to many factors, some of which include theimprovements in the equipment available for bioreactors, the increasedunderstanding of biological systems and increased knowledge as to theinteractions of materials (such as monoclonal antibodies and recombinantproteins) with the various systems of the human body.

A perfusion bioreactor processes a continuous supply of fresh media thatis fed into the bioreactor while growth-inhibiting byproducts areconstantly removed. The nonproductive downtime can be reduced oreliminated with a perfusion bioreactor process. The cell densitiesachieved in perfusion culture (30-100 million cells/mL) are typicallyhigher than for fed-batch modes (5-25 million cells/mL). Theseimprovements have led to lower contamination in the harvest and betteryields without significant increase in cost. A perfusion bioreactor mayuse a cell retention device to prevent escape of the culture whenbyproducts are being removed. These cell retention systems may add alevel of complexity to the perfusion process, where further management,control, and/or maintenance operations may be implemented.

Operational issues such as malfunction or failure of the cell retentionequipment has previously been a problem with perfusion bioreactors,which has limited their attractiveness in the past.

In the biotechnology and biopharma fields, it is desirable to separatemany different types of materials from a primary fluid stream based onsize, surface active materials, density, and other characteristics.Tangential flow filtration has been widely adopted to separate andconcentrate diverse array of microbes and other materials from a fluidstream. These materials have been filtered utilizing different physicalpolymer membranes such as polyvinylidene difluoride (PDF) and polyethersulfone (PES).

Issues that arise with tangential flow filtration include the cost ofthe membrane materials, the formation of a gel layer orfiltrate-suppressing layer on the filter membranes, the consistency ofthe polymer membranes, and the entire efficiency of the process. Thereis also the possibility of product loss due to clogging of the polymermembranes and the need to replace these clogged membranes.

There is therefore a need to improve the continuous processing ofmaterials throughout the tangential flow process, decrease the cost,increase the efficiency, and improve the overall ability of theseparation process to perform primary, secondary, tertiary and beyondseparation of materials from a primary fluid stream.

BRIEF DESCRIPTION

The present disclosure relates, in various embodiments, to acousticdevices which are used for perfusion biomanufacturing. Moreparticularly, the devices are coupled to an associated bioreactor.Within the bioreactor, biomolecules, such as recombinant proteins ormonoclonal antibodies, and other materials are produced. Materials mayinclude viruses, virus-like particles, exosomes, oncozomes, materialsfrom cell free protein synthesis, cell vesicles, proteins, monoclonalantibodies, recombinant proteins and other materials. The acousticdevice is used for separating desirable products from the bioreactorcontents on a continuous basis, and the cells or other materials arecontinuously returned to the bioreactor for further culturing.Generally, a fluid mixture containing the cells and the desired productsare passed or flowed through the acoustic device and separated thereinby multi-dimensional standing wave(s). The fluid mixture generally alsocontains other materials, such as cell debris and fines. The fluidmixture can be continuously flowed into the device, with desiredproducts being continuously removed. The acoustic perfusion devicereturns healthy viable cells to the bioreactor while desired productsare harvested and flowed downstream for further processing, e.g.,additional filtering, chromatography, etc. Additionally, the cellculture media in the bioreactor is clarified as cell fragments are alsoallowed to pass into the harvest stream and thereby out of the fluidmixture being recycled to the bioreactor. This arrangement results inlower overall cell culture media usage, corresponding to a predictedcost savings of up to $20,000 per day for large bioreactors.

Disclosed in various embodiments are acoustic perfusion devices,comprising: an acoustic chamber; an inlet port, an inlet flow pathleading from the inlet port to the acoustic chamber; an outlet port forrecirculating fluid flowing through the device back to its source (e.g.a bioreactor), and an outlet flow path leading from the acoustic chamberto the outlet port; at least one collection or harvest port forcollecting a product stream of fluid exiting the acoustic chamber; andat least one ultrasonic transducer in the acoustic chamber below the atleast one harvest port, the at least one ultrasonic transducer includinga piezoelectric material driven by a drive signal to create an acousticstanding wave across a collection or harvest flow path leading from theacoustic chamber to the at least one collection or harvest port. Theacoustic standing wave may be planar or multi-dimensional, or acombination of such waves may be present within the acoustic chamber(generally from multiple transducers). The acoustic standing wave can bethought of as a “force field” that holds back whole cells but permitssmaller materials such as the desired biomolecules (e.g. recombinantproteins and/or monoclonal antibodies) and cell fragments, to passthrough and be removed from the fluid that is returned to thebioreactor.

The outlet port is generally below the inlet port, and is generallylocated at a bottom end of the device.

As mentioned above, the device may have one or more collection orharvest ports at the top of the device. In some more specificembodiments, the device may have a total of two harvest ports spacedapart from each other on the top end of the device.

In particular embodiments, the inlet port is at a first end of thedevice at a first height, the at least one ultrasonic transducer is at asecond height above the first height, and a bottom wall extends from theinlet port to the outlet port. The outlet port may be located at asecond end of the device opposite the first end. The bottom wall may beconcave, relative to a line between the inlet port and the outlet port.The device may include an upper wall above the inlet flow path. Theinlet port, the outlet port, and the at least one harvest port aresometimes all located on a front wall of the device. The front wallitself may be planar (i.e. flat).

The device can further comprise a reflector located in the acousticchamber opposite the at least one ultrasonic transducer. Alternatively,the device can have a total of two ultrasonic transducers located onopposite sides of the harvest flow path at the same height and facingeach other, or additional ultrasonic transducers can be located onmultiple sides of the collection/harvest flow path. A reflector may belocated between the two ultrasonic transducers. There may also be aplurality of transducer/reflector pairs located as appropriate to formplanar, multi-dimensional, or combinations of such acoustic standingwave(s).

In particular embodiments, the acoustic standing wave results in anacoustic radiation force having an axial force component and a lateralforce component that are of the same order of magnitude.

In other embodiments of the device disclosed herein, the inlet flow pathleads from the inlet port downwards towards a bottom end of the deviceand past the outlet port, and then upwards to the acoustic chamber.Sometimes, the inlet port and the at least one harvest port are bothlocated on a top wall of the device, and the outlet port is located on afront wall of the device. The at least one ultrasonic transducer may bemounted in a rear wall or a front wall of the device. The bottom wall ofthis acoustic chamber can be a sloped planar surface. The reflector maybe made of a transparent material.

The inlet flow path may be shaped to generate a tangential flow pathbelow an acoustic field generated by the acoustic standing wave. Instill additional versions seen herein, the inlet flow path enters theacoustic chamber on a first side of the device, and the outlet port islocated (i) on the first side of the device or (ii) on a second oppositeside. The inlet port can be located on a front side of the device, andthe at least one harvest port can be located on a top wall of thedevice. The at least one transducer can be located on a front side or arear side of the device. In more particular embodiments, there can betwo transducers, one on the front side and one of the rear side. In yetother particular embodiments, there is an ultrasonic transducer on thefront or rear side, and a reflector located on the respective rear orfront side opposite the transducer.

In additional embodiments, the perfusion device further comprises arecirculation flow path between the inlet port and the outlet port thatdoes not enter the acoustic chamber, and the recirculation flow path islocated below the acoustic chamber. In some particular embodiments, theinlet flow path travels through a different passage than the outlet flowpath. In yet other embodiments, the inlet flow path and the outlet flowpath travel through a common passage.

The device may be attached to a mounting piece having holes forattachment.

Also disclosed are methods for separating cells from a fluid mixturecontaining the cells. The fluid mixture is flowed through an acousticperfusion device of the structure described above, having at least oneultrasonic transducer. The at least one ultrasonic transducer is drivento create the acoustic standing wave. A fluid enriched in cells can becollected from the outlet port and a clarified fluid, depleted in cells,can be collected from the at least one harvest port.

In particular embodiments, the flow rate through the collection/harvestflow path is at least one order of magnitude smaller than a flow ratethrough the inlet flow path. In specific embodiments, a flow rate of thefluid mixture entering the device through the inlet port is about 1liter per minute and a flow rate of the fluid depleted in cells exitingthe device through the at least one collection/harvest port is about 10milliliters per minute. Alternatively, the ratio of the flow rateentering through the inlet port to the flow rate exiting through the atleast one collection/harvest port is such that the acoustic standingwave is not overcome by the main body of cells, or in other words sothat a large volume of cells do not begin exiting the device through thecollection/harvest port(s).

The methods may further comprise pulling the fluid mixture through thedevice using a first pump attached to the at least one harvest port ofthe device and a second pump attached to the outlet port of the device.

Also disclosed herein are flow devices adapted to (i) receive a flowingmixture containing a primary fluid and cells; and (ii) to use a firstacoustic standing wave to continuously draw off a harvest fluid streamdepleted in cells from the flowing mixture, thereby changing the cellconcentration of the flowing mixture. A pressure rise may be generatedon the upstream interface region of the acoustic standing wave, alongwith an acoustic radiation force acting on the incoming suspendedparticles. This “interface effect”, which may also be termed “edgeeffect”, acts as a barrier and is typically located at the upstreambounding surface of the volume of fluid that is ensonified by thetransducer (i.e. the flow mixture crosses the interface region to enterthe ensonified volume of fluid). The frequency of the acoustic standingwave may be modified such that different contrast factor materials maybe held back or allowed through the acoustic standing wave, or such thatparticles of one given size range are retained and particles of a secondgiven range are allowed to flow through the standing wave.

The device may further comprise a secondary flow chamber in which theharvest fluid stream depleted in cells passes through a second acousticstanding wave having a frequency different from, or equal to the firstacoustic standing wave. For example, the second acoustic standing wavemay have a higher or lower frequency than the first acoustic standingwave. The ratio of the frequency of the two standing waves is, in someembodiments, at least 2:1 (i.e. one of the frequencies is at least twicethe other frequency, e.g. 3 MHz and 6 MHz).

Also disclosed herein are flow devices that comprise: at least one inletfor receiving a flowing mixture of a primary fluid and cells, anultrasonic transducer that produces a first ultrasonic acoustic standingwave and uses a pressure rise and an acoustic radiation force generatedon an upstream interface region of the first ultrasonic acousticstanding wave to separate the flowing mixture into a primary high cellconcentration fluid stream and a secondary harvest fluid stream; anoutlet port for the primary high cell concentration fluid stream; and atleast one collection port for the secondary harvest fluid stream. Ableed port can also be present for extracting a concentrated fluid/cellmixture. The fluid mixture may comprise particles such as mammaliancells, bacteria, cell debris, fines, proteins, exosomes, vesicles,viruses, and insect cells.

The device may further comprise a secondary flow chamber in which thesecondary harvest fluid stream passes through a second acoustic standingwave having a frequency different from, or equal to, the firstultrasonic acoustic standing wave.

Also disclosed herein are methods for separating microspheres,microbeads, nano beads, microcarriers and other micro and nanoparticulates (collectively known as beads), and cells from a host fluid.The methods comprise flowing a mixture containing a host fluid, thebeads, and the cells through an acoustophoretic device; sending avoltage signal to drive an ultrasonic transducer to create themulti-dimensional acoustic standing wave, wherein a recirculating fluidstream having a tangential flow path is located substantially tangentialto the multi-dimensional acoustic standing wave and separated therefromby an interface region; and wherein at least a portion of the cells(e.g., at least 95% of the cells, including up to about 99% of thecells) pass through the multi-dimensional acoustic standing wave, andthe beads are held back from the multi-dimensional acoustic standingwave in the recirculating fluid stream at the interface region. Theacoustophoretic device may be constructed as described herein, namelycomprising a flow chamber having at least one inlet and at least oneoutlet; at least one ultrasonic transducer located in the flow chamber,the at least one ultrasonic transducer including a piezoelectricmaterial driven by a voltage signal to create a multi-dimensionalacoustic standing wave in the flow chamber; and a reflector located inthe flow chamber opposite from the at least one ultrasonic transducer.

In particular embodiments, the beads are non-functionalized. Otherembodiments include functionalizing the beads such that the beads attachto the cells. The beads can be functionalized with various materials,such as antigens, on the surface that will allow for affinity binding ofmaterials (e.g., viruses, virus-like particles, exosomes, oncozomes,materials from cell free will protein synthesis, proteins, monoclonalantibodies, recombinant proteins) to the beads. The functionalized beadsmay, in some embodiments, have a positive contrast factor. Thefunctionalized beads having a positive contrast factor may be selectedfrom the group consisting of polystyrene beads and glass beads. In otherembodiments, have a negative contrast factor. The functionalized beadshaving a negative contrast factor may be selected from the groupconsisting of microbubbles and micro-glass spheres. The beads may, incertain embodiments, be polymeric, glass, hollow, or gas-filled. Thebeads can be spherical, toroidal, cylindrical, or conical.

In certain embodiments of the method, the mixture can includemicrovesicles (e.g., exosomes, oncosomes, viruses, proteins, recombinantproteins, and monoclonal antibodies). A pressure rise and an acousticradiation force on cells can be generated at the interface region toclarify the host fluid as it passes through the multi-dimensionalacoustic standing wave. In particular embodiments, the method furthercomprises collecting the cells after passing through themulti-dimensional acoustic standing wave for a first time andrecirculating the collected cells are back through the device forseparation by the multi-dimensional acoustic standing wave for a secondtime. The method can further comprise collecting the cells after passingthrough the multi-dimensional acoustic standing wave for the secondtime.

Also disclosed herein are acoustophoretic devices including a flowchamber having at least one inlet port and at least one outlet; at leastone ultrasonic transducer coupled to the flow chamber and at least onereflector coupled to the flow chamber opposite the at least oneultrasonic transducer, wherein the at least one ultrasonic transducerincludes a piezoelectric material driven by a voltage signal to create amulti-dimensional acoustic standing wave in the device; and a coolantinlet adapted to permit the ingress of an associated cooling fluid intothe device for cooling the at least one ultrasonic transducer.

In certain constructions, the at least one inlet port includes a firstinlet port on a front wall of the device, a second inlet port on a rearwall of the device opposite the front wall thereof, and the at least oneoutlet includes a permeate outlet on the first sidewall of the device, awastewater outlet on the first sidewall of the device, and a drainoutlet on a first sidewall of the device.

The cooling fluid can be the same or different from the primary fluidthat flows through the flow chamber carrying particles (e.g., cells andbeads). For example, the cooling fluid can be water, air, alcohol,ethanol, ammonia, or some combination thereof. The cooling fluid can, incertain embodiments, be a liquid, gas, or gel. The cooling fluid can bean electrically non-conductive fluid to prevent electric short-circuits.

These and other non-limiting characteristics are more particularlydescribed below.

BRIEF DESCRIPTION OF THE DRAWINGS

The following is a brief description of the drawings, which arepresented for the purposes of illustrating the exemplary embodimentsdisclosed herein and not for the purposes of limiting the same.

FIG. 1 illustrates a single standing acoustic wave generated by anultrasonic transducer and a reflector.

FIG. 2 is an illustration comparing a fed-batch bioreactor system with aperfusion bioreactor system.

FIG. 3 is a cross-sectional view that shows the various components of astirred-tank bioreactor.

FIG. 4 is a perspective view of one exemplary embodiment of an acousticperfusion device of the present disclosure, having two collection orharvest ports and a single ultrasonic transducer.

FIG. 5 shows a second exemplary embodiment of an acoustic perfusiondevice of the present disclosure, with a single reflector locatedbetween two ultrasonic transducers.

FIG. 6 is a schematic view illustrating a perfusion bioreactor coupledwith an acoustic perfusion device of the present disclosure, and arecycle path.

FIG. 7 is a front cross-sectional view of a third exemplary embodimentof an acoustic perfusion device of the present disclosure.

FIG. 8 is an exterior perspective view of the acoustic perfusion deviceof FIG. 7.

FIG. 9 is a front cross-sectional view of a fourth exemplary embodimentof an acoustic perfusion device of the present disclosure.

FIG. 10 is a perspective view of the acoustic perfusion device of FIG.9.

FIG. 11 is a cross-sectional diagram of a conventional ultrasonictransducer.

FIG. 12 is a cross-sectional diagram of an ultrasonic transducer of thepresent disclosure. An air gap is present within the transducer, and nobacking layer or wear plate are present.

FIG. 13 is a cross-sectional diagram of an ultrasonic transducer of thepresent disclosure. An air gap is present within the transducer, and abacking layer and wear plate are present.

FIG. 14 is a graph of electrical impedance amplitude versus frequencyfor a square transducer driven at different frequencies.

FIG. 15 illustrates the trapping line configurations for seven of theresonance frequencies (minima of electrical impedance amplitudes) ofFIG. 14 from the direction orthogonal to fluid flow.

FIG. 16 is a computer simulation of the acoustic pressure amplitude(right-hand scale in Pa) and transducer out of plane displacement(left-hand scale in meters). The text at the top of the left-hand scalereads “×10⁻⁷”. The text at the top of the left-hand scale by theupward-pointing triangle reads “1.473×10⁻⁶”. The text at the bottom ofthe left-hand scale by the downward-pointing triangle reads“1.4612×10⁻¹⁰”. The text at the top of the right-hand scale reads“×10⁶”. The text at the top of the right-hand scale by theupward-pointing triangle reads “1.1129×10⁶”. The text at the bottom ofthe right-hand scale by the downward-pointing triangle reads “7.357”.The triangles show the maximum and minimum values depicted in thisfigure for the given scale. The horizontal axis is the location withinthe chamber along the X-axis, in inches, and the vertical axis is thelocation within the chamber along the Y-axis, in inches.

FIG. 17 shows the In-Plane and Out-of-Plane displacement of a crystalwhere composite waves are present.

FIG. 18 is a view of a first acoustic perfusion device of the presentdisclosure fluidly connected to an associated bioreactor, showing aplurality of hoses fluidly connecting the various ports of the device tothe associated bioreactor and an outflow pump fluidly connecting theoutlet port of the device to the associated bioreactor.

FIG. 19 is a view of another acoustic perfusion device of FIG. 5,showing a reflector in the acoustic chamber between first and secondtransducers. A fluid mixture is also present in the device and arrowsare shown indicating the direction of flow in addition to wavesindicating the acoustic field between the reflector and first and secondtransducers.

FIG. 20 is a graph showing the efficiency of removing cells from a fluidmixture for one experiment at two different perfusate/feed rates.

FIG. 21 is a graph showing the harvest flow (also referred to as theperfusate) turbidity reduction for an experiment

FIG. 22 is a graph showing the cell viability for varied flow rates forthe experiment conducted for the graphs of FIGS. 20-21.

FIG. 23 is a graph showing the total cell density and cell retention forvaried flow rates and flow methods for another experiment.

FIG. 24 is a graph showing the cell viability for varied flow rates forthe experiment conducted for the graphs of FIG. 23.

FIG. 25 is a graph showing the total cell density and cell retention forvaried numbers of ultrasonic transducers for another experiment.

FIG. 26 is a graph showing the cell viability for varied numbers ofultrasonic transducers for the experiment conducted for the graph ofFIG. 25.

FIG. 27 is a picture of another acoustic perfusion device that wastested.

FIG. 28 is a graph showing the effect of the perfused flow rate or thetransducer voltage on the cell retention.

FIG. 29A is a microscope image from a ViCell Cell Analyzer of theparticles in the feed stream going into the device. The feed stream is abioreactor fluid containing CHO cells, protein, and cell fragments. FIG.29B is a graph of the particle diameter distribution of the feed,showing a bimodal distribution. In FIG. 29B, the y-axis is the particlecount from 0 to 200 in intervals of 10. The x-axis is the particlediameter in microns from 6 to 50 in intervals of 2. The total number ofparticles is 5539, the mean particle size is 16.78 microns, standarddeviation of 6.76 microns, and mode particle size of 10.56 microns.

FIG. 30A is a microscope image of the perfusate (or clarified harvestflow) exiting the device. FIG. 30B is a graph of the particle diameterdistribution of the perfusate, showing a unimodal distribution at muchlower sizes. In FIG. 30B, the y-axis is the particle count from 0 to 300in intervals of 20. The x-axis is the particle diameter in microns from6 to 50 in intervals of 2. The total number of particles is 2919, themean particle size is 10.08 microns, standard deviation of 3.75 microns,and mode particle size of 8.99 microns.

FIG. 31 is a CFD model showing the velocity distribution within thedevice of FIG. 27. The text at the top of the scale reads “×10⁻²”. Thetext at the top of the scale by the upward-pointing triangle reads“0.678”. The text at the bottom of the scale by the downward-pointingtriangle reads “0”. The triangles show the maximum and minimum valuesdepicted in this figure for the given scale. The scale runs from 0 to 5m/s in intervals of 0.5, with black indicating 5 at the top of thescale, and white indicating zero at the bottom of the scale.

FIG. 32 is a front view of the device of FIG. 27, showing the flowpaths, acoustic field, and acoustic interface effect.

FIG. 33 is a composite photograph showing the acoustic perfusion deviceof FIG. 27 in two operating modes. On the left, the device is in startupor cell settling mode. On the right, the device is in steady cellretention mode.

FIG. 34 shows the geometry of a model simulation of the acoustic deviceused for cell retention. The model contains two fluids, one a clarifiedfluid within the acoustic field, the other a high cell density fluid tothe left of the acoustic field, a piezoelectric transducer, a steelreflector, and an aluminum housing. The first fluid was water within theacoustic field and the second fluid was a 15% concentration of CHO cellsin water solution outside (to the left) of the acoustic field. The bluesolid line in the model indicates the separation line between the twofluids.

FIGS. 35A, 35B, and 35C are graphs showing the displacement of thepiezoelectric material, the aluminum housing, and the steel reflector(left-side scale); and the acoustic pressure in the two fluids(right-side scale) of the model simulation of FIG. 34 at severalfrequencies of operation. FIG. 35A is at a frequency of 2.218 MHz. FIG.35B is at a frequency of 2.2465 MHz. FIG. 35C is at a frequency of2.3055 MHz. For all three graphs, the left-side scale is indicated withtext at the top of the scale reading “×10⁻⁶” or “×10⁻⁷”, and is in unitsof inches. The right-side scale is indicated with text at the top of thescale reading “×10⁶”, and is in units of Pascals. The y-axis runs from−0.8 to 1.6 in intervals of 0.2. The x-axis runs from −0.5 to 1.5 inintervals of 0.5.

FIG. 36 is a graph showing the average lateral force (N) and the averagelateral force normalized by power (N/W) acting on suspended CHO cells atseveral frequencies of operation.

FIG. 37 is a picture (top view) of an acoustic perfusion device of thepresent disclosure. Arrows indicate the flow into the inlet port; theflow out of the outlet port; the clarified fluid flow out the top of thedevice and the flow of concentrate out the bottom of the device.

FIG. 38 is a picture (side view) of the acoustic perfusion device ofFIG. 37.

FIG. 39 is a graph of cell retention vs. perfusate flowrate for thedevice of FIG. 37.

FIG. 40 is a prior art illustration showing direct flow filtration (DFF)and tangential flow filtration (TFF).

FIG. 41 is a picture illustrating a first mode of operation duringperfusion, in which cells are trapped, clustered, and separated from aharvest stream. The device is operated vertically, with an arrowindicating the direction of gravity.

FIG. 42 is a picture illustrating a second mode of operation duringperfusion, in which cells are prevented from entering an acousticstanding wave field while smaller particles are permitted to passthrough the field and into the harvest stream. The device is operatedvertically, with an arrow indicating the direction of gravity.

FIG. 43 is a perspective view of a fifth exemplary embodiment of anacoustic perfusion device of the present disclosure. This embodimentincludes a direct recirculation flow path between the inlet port and theoutlet port. An inflow passageway and an outflow passageway join therecirculation flow path to the acoustic chamber, and create a tangentialsweeping flow underneath the acoustic field.

FIG. 44 is a front view picture of the device of FIG. 43. The inflowpassageway and the outflow passageway are clearly visible, along withthe recirculation pipe.

FIG. 45 is a diagrammatic side view of the device of FIG. 43.

FIG. 46A and FIG. 46B are a diagrammatic front view of the device ofFIG. 43, showing one internal structure. The figures are duplicated dueto the number of reference numerals.

FIG. 46C and FIG. 46D are a diagrammatic front view of the device ofFIG. 43, showing an alternative internal structure, where the inflowpassageway and the outflow passageway have a different flow geometry.The figures are duplicated due to the number of reference numerals.

FIG. 47 is a perspective view of a sixth exemplary embodiment of anacoustic perfusion device of the present disclosure. This embodimentincludes a direct recirculation flow path between the inlet port and theoutlet port. A single passageway joins the recirculation flow path tothe acoustic chamber, and acts as both the inflow passageway and theoutflow passageway.

FIG. 48 is a diagrammatic side view of the device of FIG. 43.

FIG. 49 is a diagrammatic front view of the device of FIG. 43, showingthe internal structure.

FIG. 50 is a graph of cell retention versus time for the device of FIG.7. The y-axis runs from 90% to 100% in intervals of 1%. The x-axis runsfrom 0 to 165 minutes in intervals of 15 minutes. Tests were performedon two different days.

FIG. 51 is a graph of perfusate cell density (million cells/mL) versustime for the device of FIG. 7. The y-axis runs from 0.00 to 3.50 inintervals of 0.50. The x-axis runs from 0 to 165 minutes in intervals of15 minutes. Tests were performed on two different days.

FIG. 52 is a CFD model showing the velocity distribution within thedevice of FIG. 46A. The text at the top of the scale reads “×10⁻³”. Thescale runs from −5 to 5 m/s in intervals of 1.

FIG. 53 is a CFD model showing the velocity distribution within thedevice of FIG. 46C at 250 mL/min flowrate. The text at the top of thescale reads “×10⁻³”. The scale runs from −5 to 5 m/s in intervals of 1.

FIG. 54 is a CFD model showing the velocity distribution within thedevice of FIG. 46C at 1000 mL/min flowrate. The text at the top of thescale reads “×10⁻³”. The scale runs from −5 to 5 m/s in intervals of 1.

FIG. 55 is a CFD model showing the velocity distribution for anotherinternal structure for the device of FIG. 43. The scale runs from −0.05to 0.1 m/s in intervals of 0.05.

FIG. 56 is a CFD model showing the velocity distribution for anotherinternal structure for the device of FIG. 43. The scale runs from −0.1to 0.15 m/s in intervals of 0.05.

FIG. 57 is a graph of cell retention versus time for the device of FIG.46A. The y-axis runs from 90% to 100% in intervals of 1%. The x-axisruns from 0 to 100 minutes in intervals of 10 minutes.

FIG. 58 is a graph of perfusate cell density (million cells/mL) versustime for the device of FIG. 46A. The y-axis runs from 0.00 to 1.20 inintervals of 0.20. The x-axis runs from 0 to 100 minutes in intervals of10 minutes.

FIG. 59 is a graph of cell retention versus time for the device of FIG.46A at two different frequencies, 1 MHz and 2 MHz. The darker circlesare for 2 MHz. The y-axis runs from 0.00 to 1.00 in intervals of 0.10.The x-axis runs from 0 to 300 minutes in intervals of 50 minutes.

FIG. 60 is a graph of perfusate cell density (million cells/mL) versustime for the device of FIG. 46A at two different frequencies, 1 MHz and2 MHz. The darker circles are for 2 MHz. The y-axis runs from 0.00 to25.00 in intervals of 5.00. The x-axis runs from 0 to 300 minutes inintervals of 50 minutes.

FIG. 61 is a front cross-sectional view of a seventh exemplaryembodiment of an acoustic perfusion device of the present disclosure.

FIG. 62 is a front view photograph of a working 1 inch by 1 inch by 1inch separator system using a 1 inch by 1 inch crystal to separate beadsfrom a yeast mixture.

FIG. 63 is a log-log graph showing the relationship of the MRC-5radiation force, the Cytodex beads radiation force, gravity/buoyancyforce, and Stokes' drag force to particle size. The horizontal axis isin microns (μm) and the vertical axis is in Newtons (N).

FIG. 64 is a graph of viable cell density (VCD) (10⁶ cells/mL) over timefor specified AWS power (%) and specified harvest to recirculationratios (%). The x-axis (culture day) runs from 0 to 30 days in intervalsof 5 days. The left-side y-axis (VCD) runs from 0 to 80×10⁶ cells/mL inintervals of 1 before the break and intervals of 20 after the break, andis used for the circles. The first right-side y-axis (AWS power) runsfrom 0 to 60 in intervals of 10, and applies to the line that starts atabout 30% and ends at about 10%. The second right-side y-axis (harvestto recirculation ratio) runs from 0.0 to 0.5 in intervals of 0.1, andapplies to the line that starts at about 0.15% and ends at 0.3%. Avertical dashed line shows that AWS began on day 7. The darker lineextending from the vertical dash line corresponds to the AWS power, andthe lighter line extending from the vertical dashed line corresponds tothe harvest to recirculation ratio. The white dots represent the AWSharvest, and the black dots represent the bioreactor.

FIG. 65 is a space characterization of the harvest to recirculation flowratio (%) versus the AWS power (%). The y-axis (harvest to recirculationflow ratio) runs from 0.2 to 1.2 in intervals of 0.2. The x-axis (AWSpower) runs from 10 to 60 in intervals of 10.

FIGS. 66A-66C are graphs of AWS versus TFF mAb titer and producttransmission. In all three graphs, AWS is diamonds, TFF with bleed issquares, and TFF without bleed is triangles. FIG. 66A shows the mAbtiter on the y-axis versus the culture day on the x-axis. The productionof the mAb titer is higher for the AWS than it is for the TFF where thecells are bled away. FIG. 66B is a graph comparing the transmission ofmonoclonal antibody (mAb) versus culture day. The y-axis is %transmission, and the x-axis is the culture day. The dotted lineindicates 100% transmission. Higher % transmission is better. The mAbtransmission is higher for the AWS versus the TFF with or without cellbleed. FIG. 66C shows relative mAb residence time on the y-axis versusculture day on the x-axis. The relative residence time for the mAbseparated by AWS is lower than the relative residence time for the TFFwith or without cell bleed.

FIG. 67 is a graph of cell separation efficiency over time for AWS cellseparation at a 10 L scale. The x-axis (culture day) runs from 0 to 16days in intervals of 2 days, and is used for the middle line that startsat x=0 days. The first left-side y-axis (harvest VCD) runs from 0 to8×10⁶ cells/mL in intervals of 2, and is used for the bottom line thatstarts at roughly x=1 day. The second left-side y-axis (bioreactor VCD)runs from 0 to 80×10⁶ cells/mL in intervals of 20. The right-side y-axis(cell separation efficiency) runs from 0.0 to 1.0 in intervals of 0.2,and is used for the top line. The cell separation efficiency is apercentage (i.e., 0.2 corresponds to 20% efficiency, 0.6 corresponds to60% efficiency, 1.0 corresponds to 100% efficiency, etc.)

FIG. 68A is a graph showing the relative abundance of high molecularweight (HMW) species, monomer species, and low molecular weight (LMW)species of a monoclonal antibody (mAb) after size exclusionchromatography (SEC). The darker bars represent the bioreactor, and thelighter bars represent the harvest. FIG. 68B is a graph showing therelative abundance of acidic, main peak, and basic species of a mAbafter ion exchange chromatography (IEX). Again, the darker barsrepresent the bioreactor, and the lighter bars represent the harvest.FIG. 68C is a bar graph showing glycosylation of proteins, showing theproteins sent through the acoustic perfusion device do not differ fromproteins sent through a TFF device. BRX refers to the bioreactor, andHARV indicates the harvested material.

FIGS. 69A-69B are graphs comparing the AWS process to a tangential flowfiltration (TFF) perfusion process using a 0.22 micron membrane. Thedarker bars represent AWS and the lighter bars represent TFF. FIG. 69Ashows the relative abundance of HMW species, monomer species, and LMWspecies of a mAb after size exclusion chromatography (SEC). FIG. 69Bshows the relative abundance of acidic, main peak, and basic species ofa mAb after IEX.

FIG. 70 is a perspective view of an eighth exemplary embodiment of anacoustic perfusion device of the present disclosure.

DETAILED DESCRIPTION

The present disclosure may be understood more readily by reference tothe following detailed description of desired embodiments and theexamples included therein. In the following specification and the claimswhich follow, reference will be made to a number of terms which shall bedefined to have the following meanings.

Although specific terms are used in the following description for thesake of clarity, these terms are intended to refer to the particularstructure of the embodiments selected for illustration in the drawings,and are not intended to define or limit the scope of the disclosure. Inthe drawings and the following description below, it is to be understoodthat like numeric designations refer to components of like function.

The singular forms “a,” “an,” and “the” include plural referents unlessthe context clearly dictates otherwise.

The term “comprising” is used herein as requiring the presence of thenamed component and allowing the presence of other components. The term“comprising” should be construed to include the term “consisting of”,which allows the presence of only the named component, along with anyimpurities that might result from the manufacture of the namedcomponent.

Numerical values should be understood to include numerical values whichare the same when reduced to the same number of significant figures andnumerical values which differ from the stated value by less than theexperimental error of conventional measurement technique of the typedescribed in the present application to determine the value.

All ranges disclosed herein are inclusive of the recited endpoint andindependently combinable (for example, the range of “from 2 grams to 10grams” is inclusive of the endpoints, 2 grams and 10 grams, and all theintermediate values). The endpoints of the ranges and any valuesdisclosed herein are not limited to the precise range or value; they aresufficiently imprecise to include values approximating these rangesand/or values.

The modifier “about” used in connection with a quantity is inclusive ofthe stated value and has the meaning dictated by the context. When usedin the context of a range, the modifier “about” should also beconsidered as disclosing the range defined by the absolute values of thetwo endpoints. For example, the range of “from about 2 to about 10” alsodiscloses the range “from 2 to 10.” The term “about” may refer to plusor minus 10% of the indicated number. For example, “about 10%” mayindicate a range of 9% to 11%, and “about 1” may mean from 0.9-1.1.

It should be noted that many of the terms used herein are relativeterms. For example, the terms “upper” and “lower” are relative to eachother in location, i.e. an upper component is located at a higherelevation than a lower component in a given orientation, but these termscan change if the device is flipped. The terms “inlet” and “outlet” arerelative to a fluid flowing through them with respect to a givenstructure, e.g. a fluid flows through the inlet into the structure andflows through the outlet out of the structure. The terms “upstream” and“downstream” are relative to the direction in which a fluid flowsthrough various components, i.e. the flow fluids through an upstreamcomponent prior to flowing through the downstream component. It shouldbe noted that in a loop, a first component can be described as beingboth upstream of and downstream of a second component.

The terms “horizontal” and “vertical” are used to indicate directionrelative to an absolute reference, i.e. ground level. However, theseterms should not be construed to require structures to be absolutelyparallel or absolutely perpendicular to each other. For example, a firstvertical structure and a second vertical structure are not necessarilyparallel to each other. The terms “top” and “bottom” or “base” are usedto refer to surfaces where the top is always higher than the bottom/baserelative to an absolute reference, i.e. the surface of the earth. Theterms “upwards” and “downwards” are also relative to an absolutereference; upwards is always against the gravity of the earth.

The present application refers to “the same order of magnitude.” Twonumbers are of the same order of magnitude if the quotient of the largernumber divided by the smaller number is a value of at least 1 and lessthan 10.

The term “beads” is used herein to refer to microspheres, microbeads,nano beads, microcarriers and other micro and nano particulates that arein the micrometer and nanometer size range, which may be functionalizedor non-functionalized, and are composed of any one or more of a varietyof materials and may be of one or more various shapes.

Bioreactors are useful for making biomolecules such as recombinantproteins or monoclonal antibodies. Very generally, cells are cultured ina bioreactor vessel with media in order to produce the desired product,and the desired product is then harvested by separation from the cellsand media in an acoustic perfusion device, such as the device of thepresent disclosure. The acoustic filtering device permits the withdrawalof some desired product, a small portion of the media, and cellularfragments/debris smaller than the cells, with the remainder beingrecycled back to the bioreactor (particularly the cells). The use ofmammalian cell cultures including Chinese hamster ovary (CHO), NSOhybridoma cells, baby hamster kidney (BHK) cells, insect cells, andhuman cells (e.g. T-cells, B-cells, stem cells, red blood cells), andliving/biological cells in general has proven to be a very efficaciousway of producing/expressing the recombinant proteins and monoclonalantibodies used in various applications such as pharmaceuticals orvaccines. Two general types of bioreactor processes exist: fed-batch andperfusion.

While fed-batch reactors are the norm currently, due mainly to thefamiliarity of the process to many scientists and technicians, perfusiontechnology is growing at a very fast rate. Many factors favor the use ofa perfusion bioreactor process, primarily because it is conducive tocontinuous production. The capital and start-up costs for perfusionbioreactors are lower, there is a smaller demand on upstream anddownstream capacity, throughput can be higher, the process iscontinuous, and the process uses smaller volumes and fewer seed stepsthan fed-batch methods. A perfusion bioreactor process also lends itselfbetter to development, scale-up, optimization, parameter sensitivitystudies, and validation.

A perfusion bioreactor may also be utilized to generate cells that wouldbe utilized in a cell therapy process. In this type of perfusionbioreactor, biological cells such as CAR T-cells, Jurkat T-cells and thelike are cultured in a perfusion bioreactor. The acoustic standing waveused in the perfusion devices of the present disclosure can be used toseparate viable and nonviable cells after the transfection process. Thisseparation ability allows for improved efficacy of the inoculation ofthe patient with this T-cell therapy, since viable cells may beprimarily utilized. The nonviable cells and cell fragments are separatedout through the perfusion process, with these materials going into thesecondary flow and exiting the bioreactor.

A perfusion bioreactor may also be used for production of exosomes,microvesicles, or vesicles by cells. The acoustic perfusion device canthen be used to harvest the exosomes, or other desired cell products, orother targets in a host fluid. In a similar fashion, a perfusionbioreactor can be used to produce viruses, such as lentivirus, which areused in cell and gene therapy to transfect cells. The acoustic perfusiondevice can then be used to harvest the virus. In all cases, the deviceis a cell retention device.

Recent developments in perfusion bioreactor technology also favor itsuse. Control technology and general support equipment is improving forperfusion bioreactors, increasing the robustness of perfusion processes.The perfusion process can now be scaled up to bioreactors having avolume up to 1000 liters (L). Better cell retention systems forperfusion bioreactors result in lower cell loss and greater celldensities than have been seen previously. Cell densities greater than 50million cells/mL are now achievable, compared to fed-batch celldensities of around 20 million cells/mL. Lower contamination andinfection rates have improved the output of perfusion bioreactors.Higher product concentrations in the harvest and better yields withoutsignificant increase in cost have thus resulted for perfusion processes.

Perfusion bioreactors are particularly attractive because of thecontinuous production of the biomolecules from the expressing cellculture, and shorter residence time of said biomolecules in the processprior to harvest. The target cells are held back by a filtrationprocess, such as tangential flow filtration (TFF) or alternatingtangential flow filtration (ATF) while the expressed biomolecules areextracted from the perfusion bioreactor. The cells are returned to thebioreactor to receive the nutrition and oxygen to maintain theproduction of the overall cell culture. In the perfusion reactorprocess, the cells continue to multiply, and some cells may be removedor bled off from the cell culture population throughout the perfusionproduction process.

The TFF and ATF processes of filtration have several issues, such asclogging/fouling and loss of biomolecule product (particularly at highcell densities), all directly related to the nature of the hollow fibermembranes used in the filtration. It is therefore desirable to find anew filtration process that does not clog and minimizes loss of thedesired biomolecule product. In addition, TFF and ATF will retain allcellular debris and fines within the bioreactor, which is not desirable.A process capable of distinguishing between cell retention whileallowing for the passing of cell debris and fines may therefore befavorable.

Briefly, the present disclosure relates to acoustic perfusion devicescapable of generating multi-dimensional acoustic standing wave(s) fromone or more piezoelectric transducers, where the transducers areelectrically excited such that they move in a multi-mode displacementpattern rather than a “piston” mode of vibration. Through this manner ofacoustic standing wave generation, a higher lateral trapping force isgenerated than if the piezoelectric transducer is excited in a “piston”mode where one large standing wave is generated. Thus, with the sameinput power to a piezoelectric transducer, the multi-dimensionalacoustic standing waves can have a higher lateral trapping forcecompared to a planar acoustic standing wave. The input power is tunablefor a controlled flow. This tunability can be used to facilitateproteinaceous fluid purification of a fluid stream coming from abioreactor. Alternatively, the acoustic standing wave may also be aplanar standing wave where the piezoelectric transducer is excited inthe piston mode, generating a planar wave. The acoustic standing wave(s)may also be a combination of planar and multi-dimensional acousticstanding waves. All of these standing waves generate an “interfaceeffect” such that the cells from the bioreactor are held back and thebiomolecule product expressed from the cells, cell fragments and smalldebris are allowed to pass through.

Acoustophoresis is a low-power, no-pressure-drop, no-clog, solid-stateapproach to particle separation from fluid dispersions (i.e., it is usedto achieve separations that are more typically performed with porousfilters, but it has none of the disadvantages of filters). Inparticular, the acoustic perfusion devices of the present disclosure aresuitable for use with macro-scale bioreactors for separations in flowingsystems with high flow rates. The acoustic perfusion device is designedto create a high intensity multi-dimensional ultrasonic standing wavethat results in an acoustic radiation force that can overcome thecombined effects of fluid drag and buoyancy or gravity at certain flowrates. As a result, the radiation force acts as a filter that preventstargeted particles (e.g., biological cells) from crossing through thestanding wave. As explained above, the trapping capability of a standingwave may be varied as desired, for example by varying the flow rate ofthe fluid, the acoustic radiation force, and the shape of the acousticfiltering device to maximize cell retention through trapping andsettling. This technology offers a green and sustainable alternative forseparation of secondary phases with a significant reduction in cost ofenergy. Excellent particle separation efficiencies have beendemonstrated for particle sizes as small as one micron.

Generally, an acoustic standing wave generates pressure minima atlocations on the standing wave where the amplitude is minimum andmaximum. These are called, respectively, nodes and anti-nodes. Thesepressure minima nodes and anti-nodes may be utilized to capturematerials that are differentiated from the surrounding environment bysize, density and compressibility (i.e., the speed of sound through thematerial). Those materials that collect at the pressure minima nodes areknown as having a positive contrast factor. Those materials that collectat the pressure minima anti-nodes are known as having a negativecontrast factor.

In a typical experiment, the system is driven such that the particlesare trapped in the ultrasonic standing wave, i.e., remain in astationary position. The axial component of the acoustic radiation forcedrives the particles, with a positive contrast factor, to the pressurenodal planes, whereas particles with a negative contrast factor aredriven to the pressure anti-nodal planes. The radial or lateralcomponent of the acoustic radiation force is the force that contributesto trapping the particle. The forces acting on the particle may begreater than the combined effect of fluid drag force and gravitationalforce.

Generally, the scattering of the acoustic field off the particlesresults in a three-dimensional acoustic radiation force, which acts as athree-dimensional trapping field. The acoustic radiation force isproportional to the particle volume (e.g., the cube of the radius) whenthe particle is small relative to the wavelength. The force isproportional to frequency and the acoustic contrast factor. The forcealso scales with acoustic energy (e.g., the square of the acousticpressure amplitude). For harmonic excitation, the sinusoidal spatialvariation of the force is what drives the particles to the stablepositions within the standing waves. When the acoustic radiation forceexerted on the particles is stronger than the combined effect of fluiddrag force and buoyancy/gravitational force, the particle is trappedwithin the acoustic standing wave field. The action of the lateral andaxial acoustic forces on the trapped particles results in formation oftightly packed clusters through concentration, clustering, clumping,agglomeration and/or coalescence of particles that, when reaching acritical size, settle continuously through enhanced gravity forparticles heavier than the host fluid or rise out through enhancedbuoyancy for particles lighter than the host fluid. Additionally,secondary inter-particle forces, such as Bjerkness forces, aid inparticle agglomeration. Relatively large solids of one material can thusbe separated from smaller particles of a different material, the samematerial, and/or the host fluid through enhanced gravitational/buoyancyseparation.

Most biological cell types present a higher density and lowercompressibility than the medium in which they are suspended, so that theacoustic contrast factor between the cells and the medium has a positivevalue. As a result, the axial acoustic radiation force (ARF) drives thecells towards the standing wave pressure nodes. The axial component ofthe acoustic radiation force drives the cells, with a positive contrastfactor, to the pressure nodes, whereas cells or other particles with anegative contrast factor are driven to the pressure anti-nodes. Theradial or lateral component of the acoustic radiation force is the forcethat traps the cells. The radial or lateral component of the ARF islarger than the combined effect of fluid drag force and gravitationalforce. For small particles or emulsions, the drag force F_(D) andStokes' drag F_(SD) can be expressed as:

$\begin{matrix}{{{\overset{harpoonup}{F}}_{D} = {4{\pi\mu}_{f}{{R_{P}( {{\overset{\_}{U}}_{f} - {\overset{\_}{U}}_{p}} )}\lbrack \frac{1 + {\frac{3}{2}\hat{\mu}}}{1 + \hat{\mu}} \rbrack}}},{{\overset{harpoonup}{F}}_{SD} = {6{\pi\mu}_{f}R_{P}\overset{arrow}{V_{s}}}}} & (1)\end{matrix}$

where U_(f) and U_(p) are the fluid and particle velocity, R_(p) is theparticle radius, μ_(f) and μ_(p) are the dynamic viscosity of the fluidand particle, respectively, {circumflex over (μ)}=μ_(p)/μ_(f) is theratio of dynamic viscosities, and {right arrow over (V_(s))} is theparticle settling velocity given by:

$\begin{matrix}{V_{S} = \frac{2( {\rho_{p} - \rho_{f}} ){gR}_{P}^{2}}{9\mu_{f}}} & (2)\end{matrix}$

where ρ_(f) is the fluid density, ρ_(p) is the particle density, and gis the universal gravitational constant.

The gravity/buoyancy force F_(B) is expressed as:

$\begin{matrix}{F_{B} = {\frac{4}{3}\pi \; R_{P}^{3}{g( {\rho_{f} - \rho_{p}} )}}} & (3)\end{matrix}$

For a particle to be trapped in the ultrasonic standing wave and developinto a particle cluster, the force balance on the particle can beassumed to be zero, and therefore an expression for lateral acousticradiation force F_(LRF) can be found, which is given by:

F _(LRF) =F _(D) +F _(B)  (4)

For a particle of known size and material property, and for a given flowrate, this equation can be used to estimate the magnitude of the lateralacoustic radiation force.

The theoretical model that is used to calculate the acoustic radiationforce is the formulation developed by Gor'kov, where the primaryacoustic radiation force F_(A) is defined as a function of a fieldpotential U, F_(A)=−∇(U), where the field potential U is defined as

$\begin{matrix}{U = {V_{O}\lbrack {{\frac{\langle{p^{2}( {x,y,z} )}\rangle}{2\rho_{f}c_{f}^{2}}f_{1}} - {\frac{3\rho_{f}{\langle{v^{2}( {x,y,z} )}\rangle}}{4}f_{2}}} \rbrack}} & (5)\end{matrix}$

and f₁ and f₂ are the monopole and dipole contributions defined by

$\begin{matrix}{{f_{1} = {{1 - {\frac{1}{{\Lambda\sigma}^{2}}\mspace{31mu} f_{2}}} = \frac{2( {\Lambda - 1} )}{{2\Lambda} + 1}}}{where}} & (6) \\{\sigma = {{\frac{c_{p}}{c_{f}}\mspace{31mu} \Lambda} = {{\frac{p_{p}}{p_{f}}\mspace{31mu} \beta_{f}} = \frac{1}{\rho_{f}c_{f}^{2}}}}} & (7)\end{matrix}$

where p is the acoustic pressure, u is the fluid particle velocity, Λ isthe ratio of cell density ρ_(p) to fluid density ρ_(f), σ is the ratioof cell sound speed c_(p) to fluid sound speed c_(f), V_(o)=πR_(p) ³ isthe volume of the cell, and < > indicates time averaging over the periodof the wave.

For a one dimensional standing wave, the acoustic pressure is expressedas

p=A cos(kx)cos(ωt)  (8)

where A is the acoustic pressure amplitude, k is the wavenumber, and ωis the angular frequency. In this case, the axial component of theacoustic radiation force F_(ARF), is found to be

$\begin{matrix}{F_{ARF} = {V_{O}{kX}\frac{A^{2}}{4\rho_{f}c_{f}^{2}}{\sin ( {2\; {kx}} )}}} & (9)\end{matrix}$

where X is the contrast factor given by

$\begin{matrix}{X = ( {\frac{{5\Lambda} - 2}{1 + {2\Lambda}} - \frac{1}{\sigma^{2}\Lambda}} )} & (10)\end{matrix}$

Particles with a positive contrast factor will be driven to the pressurenodal planes, and particles with a negative contrast factor will bedriven to the pressure anti-nodal planes. In this way, the generation ofa multi-dimensional acoustic standing wave in a flow chamber results inthe creation of tightly packed clusters of particles in the flowchamber, typically corresponding to the location of the pressure nodesor anti-nodes in the standing wave depending on acoustic contrastfactor.

Gork'ov's model is for a single particle in a standing wave and islimited to particle sizes that are small with respect to the wavelengthof the sound fields in the fluid and the particle. It also does not takeinto account the effect of viscosity of the fluid and the particle onthe radiation force. As a result, this model cannot be used for themacro-scale ultrasonic separators discussed herein since particleclusters can grow quite large. A more complex and complete model foracoustic radiation forces without any restriction as to particle sizerelative to wavelength was therefore used. The models that wereimplemented are based on the theoretical work of Yurii Ilinskii andEvgenia Zabolotskaya as described in AIP Conference Proceedings, Vol.1474-1, pp. 255-258 (2012) and “Acoustic radiation force of a spherewithout restriction to axisymmetric fields,” Proceedings of Meetings onAcoustics, Vol. 19, 045004 (2013). These models also include the effectof fluid and particle viscosity, and therefore are a more accuratecalculation of the acoustic radiation force. With these models, theacoustic radiation force is defined by

$\begin{matrix}{{\overset{arrow}{F}}_{Az} = {{\frac{i}{4}{Kk}^{2}{\sum\limits_{n = 0}^{n_{\max}}\; {\sum\limits_{m = {- n}}^{n}\; {{\sqrt{\frac{( {n + m + 1} )( {n - m + 1} )}{( {{2\; n} + 1} )( {{2\; n} + 3} )}}\lbrack {A_{n}^{*} + A_{n + 1} + {2\; A_{n}^{*}A_{n + 1}}} \rbrack}a_{n,m}^{*}a_{({{n + 1},m})}}}}} + {c.c.}}} & (11)\end{matrix}$

where a*_(n,m) and a_((n+1,m)) are the expansion coefficients ofexternal field potential with respect to spherical functionsj_(n)(kr)Y_(n,m)(θ,φ), Y_(n,m)(θ,φ) are spherical harmonies, A_(n) isthe scattering coefficient for mode (m,n), K is bulk modulus, and k isthe wavenumber in the liquid.

Cells that have a low contrast factor compared to the fluid in whichthey are transported are more difficult to separate using an acousticstanding wave. Through specialized perturbations of a piezoelectricmaterial, higher order modes of vibration in the piezoelectric materialmay be generated. When this piezoelectric material that is perturbed ina multimode fashion is coupled with a reflector, a specialized type ofacoustic standing wave, known as a multi-dimensional acoustic standingwave, is generated. In this way, target biological cells having low cellconcentrations (e.g., T cells) may be separated from a fluid mediumutilizing a multi-dimensional acoustic standing wave. The targetbiological cells are generally at lower concentrations than, forexample, a CHO cell population with 30 million cells per mL versus aconcentration of 1 million cells per mL for Jurkat T cells. Thus, thelow contrast cells, such as Jurkat T cells, in a low populationconcentration are separated continuously from the fluid media withinwhich they are entrained by utilizing a multi-dimensional acousticstanding wave.

Desirably, the ultrasonic transducer(s) generates a multi-dimensionalstanding wave in the fluid that exerts a lateral force on the suspendedparticles to accompany the axial force. The multi-dimensional standingwave generates acoustic radiation forces in both the axial direction(e.g., in the direction of the standing wave, between the transducer andthe reflector, which may be at an angle across the flow direction, andin some instances may be perpendicular to the flow direction) and thelateral direction (e.g., in the flow direction or transverse to thedirection between the transducer and the reflector). As the mixtureflows through the acoustic chamber, particles in suspension experience astrong axial force component in the direction of the standing wave.Since this acoustic force is across (e.g. perpendicular to) the flowdirection and the drag force, it quickly moves the particles to pressurenodal planes or anti-nodal planes, depending on the contrast factor ofthe particle. The lateral acoustic radiation force acts to move theconcentrated particles towards the center of each planar node, resultingin clustering, agglomeration or clumping. The lateral acoustic radiationforce component can overcome fluid drag for such clumps of particles, tocontinually grow the clusters, which can exit the mixture due to gravityor buoyancy. The drop in drag per particle as the particle clusterincreases in size, as well as the drop in acoustic radiation force perparticle as the particle cluster grows in size, may separately orcollectively influence operation of the acoustic separator device. Inthe present disclosure, the lateral force component and the axial forcecomponent of the multi-dimensional acoustic standing wave are of thesame or different order of magnitude. In some particular embodiments,the ratio of the lateral force component to the axial force component isabout 0.5 or less. In this regard, it is noted that in amulti-dimensional acoustic standing wave generated by a singletransducer, the axial force is stronger than the lateral force, but thelateral force of such a multi-dimensional acoustic standing wave is muchhigher than the lateral force of a planar standing wave, usually by twoorders of magnitude or more. Typical results published in literaturestate that the lateral force is two orders of magnitude smaller than theaxial force. In contrast, the technology disclosed in this applicationprovides for a lateral force to be of the same order of magnitude as theaxial force. However, in certain embodiments described further herein,the device use both transducers that produce multi-dimensional acousticstanding waves and transducers that produce planar acoustic standingwaves. For purposes of this disclosure, a standing wave where thelateral force is not the same order of magnitude as the axial force isconsidered a “planar acoustic standing wave.” The lateral forcecomponent of the total acoustic radiation force (ARF) generated by theultrasonic transducer(s) of the present disclosure is significant and issufficient to overcome the fluid drag force at linear velocities of upto 1 cm/s, and to create tightly packed clusters, and is of the sameorder of magnitude as the axial force component of the total acousticradiation force.

The average size of microvesicles, such as exosomes (i.e., about 150nanometers in diameter or smaller), makes it difficult to separate thesame using normal acoustophoretic techniques. Through the use of beads,the acoustic footprint or contrast factor of the exosomes may beincreased dramatically by functionalizing the surface of the bead suchthat it will attach to the exosomes and create a moiety that is subjectto acoustophoresis through its high contrast factor. The beads may haveeither a positive or negative contrast factor. Positive contrast factormaterials include polystyrene beads, glass beads and other materialswhose density is greater than that of the host fluid that contains theexosomes. Negative contrast factor materials of interest includemicrobubbles, micro-glass spheres and other low-density beads that maybe functionalized to attract exosomes and thus increase their contrastfactor footprint. The functionalization of the beads may be by, forexample, attachment of CD 9, CD 63, CD 81, or EP CAM.

The acoustophoresis separation processes of the present disclosure maybe coupled with other techniques, such as field flow fractionation wherethe ultrasonic wave is utilized to fractionate the materials entrainedin a fluid stream, such as the exosome moieties or the beads that arecarrying the exosomes through affinity binding. The acoustophoresisseparation processes of the present disclosure may also be coupled withtangential flow filtration in a flow field that is bounded by theacoustic standing wave.

In the case of exosomal separation, an efficient and cost effectiveplatform for complete isolation and segregation of extracellular vesiclepopulations from cell culturing, without the potential for lysis tooccur, is possible with the acoustic tangential flow filtrationprocesses described herein, whereas the detrimental processes of ruptureand deactivation would happen with physical filtration and/orcentrifugation. With emphasis on yielding pure exosome or oncosomepopulations that are morphologically and functionally intact, theacoustic tangential flow filtration process offers a great improvement.The acoustic tangential flow filtration process technology is readilyadapted to enable automated workflows and reduce human intervention,enable closed loop systems, and to obtain the enriched distinctpreparations of exosomes and oncosomes used to provide diagnosticdeterminations and selective therapeutic product.

It may be helpful to contrast the technology of the present disclosurewith that of prior filtration technology. FIG. 40 shows two prior artfiltration methods. The left-hand side of FIG. 40 illustrates directflow filtration (DFF). In DFF, the entire feed stream 4005 of fluid andparticles is directed towards the filter. The filter 4010 holds back theparticles 4020 that are larger than the filter's pore size, whereassmaller particles 4030 and the fluid pass through the filter. Theright-hand side of FIG. 40 illustrates tangential flow filtration (TFF).In TFF, the feed stream is not directed towards the filter. Rather, thefeed stream is directed tangentially to the filter, such that a majorityof the feed stream passes tangentially over the filter surface.Typically, this feed stream is recirculated to pass by the filter morethan once. A much smaller filtrate stream 4006 is pulled through thefilter membrane containing the smaller particles 4030. One advantage ofTFF over DFF is that the tangential stream reduces the clogging andfouling of the filter and the formation of a gel layer that sits on topof the filter.

In the devices of the present disclosure, during startup, the fluidensonified by the acoustic standing wave is clarified by the process oftrapping cells and growing them into tightly packed clusters, such thatcontinuous gravitational separation of the clusters of cells takesplace. Since there is a limited amount of new cells flowing into thisvolume, this results in a rapid clarification of the fluid subjected tothe acoustic standing wave. When this state is reached, the system canbe described as including two fluids: a first fluid containing thedesired product and small cell fragments/debris (which have passedthrough the acoustic standing wave), and a second fluid containing thebioreactor fluid and all of the cells (which are held back by theacoustic standing wave). The two fluids may be of different effectiveacoustic properties, such as density and speed of sound, with awell-defined interface between these two fluids. The acoustic standingwave is a three-dimensional acoustic field, which, in the case ofexcitation by a rectangular transducer, can be described as occupying aroughly rectangular prism volume of fluid. Typically, two opposing facesare the transducer and reflector, an adjacent pair of opposing faces arethe walls of the device, and the final opposing pair of faces, theupstream and downstream faces of the cube, extend through the fluid. Theinterface between the two fluids is generally located near the upstreamface of the acoustic standing wave field, generating an “acousticbarrier or edge effect”. This location is also referred to as anupstream interface region. The first fluid (i.e., the fluid that hasbeen clarified and contains the product, some cells, and cell fragments)is downstream of the interface and represents the harvest flow andoccupies the volume of fluid ensonified by the acoustic standing wavefield. The second fluid (i.e., the fluid containing the bioreactor fluidand most of the cells) is upstream of the interface. These two differentfluids can be seen in the photo on the right in FIG. 33. Duringoperation at increased flow rates, the interface effect location maymove downstream and is then located within the volume of fluidensonified by the transducer.

The acoustic standing wave field exerts an acoustic radiation pressure(i.e. a pressure rise) and an acoustic radiation force on the cells atthe interface region between the two fluids, thereby keeping theupstream cells from entering the acoustic field. The occurrence of theradiation pressure and the force on the interface allows for the firstfluid containing the product to pass through the interface whileretaining the cells in the upstream fluid. The cells that are held backby the effect of the acoustic radiation force at the interface betweenthe two fluids can be continuously returned to the bioreactor to ensurethey receive the nutrition and oxygen to maintain the production of theoverall cell culture.

The circulating motion of the flow field underneath the interfacetransports the cells that are retained by the acoustic field back to thebioreactor. The circulating flow motion is driven by the primaryrecirculation stream and can be optimized with acoustic chamber geometryvariations for maximum system efficiency. This process or configurationis discussed further below with respect to FIG. 33.

During perfusion, the acoustic perfusion devices of the presentdisclosure have multiple possible modes of operation. One of these modesmay be dominant in the device or they may occur concurrently dependingon the distribution of cells and fluid within the device. In a firstmode of operation illustrated in FIG. 41 (Mode 1), the fluid containinglarger particles 4020 (light color) enters the acoustic standing wavefield 4040, which is produced here between transducer 4008 and reflector4009. A multi-dimensional acoustic standing wave traps the particles atspecific points, packs the particles into tightly packed clusters 4022,and continuously separates the clusters through enhanced gravitationalsettling. The particle clusters settle out, enter the tangential flowpath (indicated by arrow 4001) and are redirected to the bioreactor bythe recirculation stream. Smaller particles 4030 (darker color) are nottrapped by, and pass through, the acoustic standing wave, to beharvested. The harvest flow direction is indicated by arrow 4002. Theorientation of this device is significant, and the direction of gravityis also indicated. In particular embodiments, the larger particles 4020can be cells and the smaller particles 4030 can be cell debris and/orfines. In other embodiments, the larger particles 4020 can be beads andthe smaller particles 4030 can be cells or exosomes (or more generallythe desired target material). In other embodiments, the larger particles4020 can be beads with attached antibodies and the smaller particles4030 can be cells that are not targeted by the functionalized beads.

The second mode of operation (Mode 2) is illustrated in FIG. 42, wherethe acoustophoretic system creates a strong barrier for cells at theinterface between the two fluids and prevents cells from entering theacoustic field. Here, a barrier of cells is established between the twofluids through the interface effect of the acoustic standing wave. Afirst clarified fluid stream 4050 contains the smaller particles/desiredbyproducts 4030 within the acoustic standing wave field and the harveststream. A second fluid stream 4055 contains the retained cells 4020upstream of the acoustic standing wave field. The harvest flow directionis indicated by arrow 4002. In this mode of operation, an acousticinterface effect is realized, as indicated by dotted line 4007(representing the interface region between the two fluids, clarifiedfluid downstream and flow mixture and cells on the upstream side). Verygenerally, the acoustic interface effect holds the cells back andprevents them from entering the acoustic field while a portion of thefluid stream containing the produced biomolecules and cell fragments ispermitted to pass through this barrier. The tangential flow pathunderneath the acoustic interface (arrow 4001) collects the retainedcells and flows them back into the main recirculation stream and back tothe bioreactor. This recirculation is discussed further below withrespect to FIG. 32 and FIG. 61. Again, the direction of gravity isindicated in the figures. Alternatively, the second fluid stream 4055can contain desired target material which is retained within theperfusion device, and the first fluid stream 4050 can be returned to thebioreactor.

In perfusion applications, the setup of the acoustophoretic device issimilar to that of TFF. A feed stream containing the cells, cell debris,fines, and product, i.e., protein, flows from the bioreactor into theperfusion system. A portion of the stream flows in a tangential fashionalong the upstream/lower interface region of the acoustic standing waveand is recirculated back to the bioreactor. A smaller portion of thefeed stream is harvested, i.e., diverted and flows through the acousticstanding wave. Here the acoustic standing wave functions very similarlyto the filter in TFF, preventing the cells from entering the acousticfield. The harvest stream contains smaller particles such as cell debrisand fines as well as the desired biomolecule product. The cells that areretained by the acoustic standing wave are transported by therecirculation stream back to the bioreactor. FIG. 32 and FIG. 61, whichare discussed further herein, also illustrate perfusion devices that usea tangential flow stream.

Perfusion applications typically entail high cell densities, e.g., >50million cells/mL, and lower harvest velocities contrary to cellclarification or oil/water applications. The two fluid streams also havedifferent effective acoustic properties, i.e., speed of sound anddensity of the media/cell mixture. As cell density increases, thedifference in acoustic properties of the two fluid streams are morepronounced. The acoustic standing wave field exerts an acousticradiation pressure, i.e., a pressure rise, on the second fluid stream,enriched with cells, as well as acoustic radiation forces on the cellssuspended in the fluid. This radiation pressure and radiation force actat the interface between the two fluids which coincides with theupstream bounding surface of the acoustic field. When this “acousticinterface” effect of acoustic radiation force is sufficiently strong, itcan prevent the cells from entering the acoustic field. Equallyimportant is a tangential flow path to collect the retained cells andtransport them back to the bioreactor.

The acoustic interface effect may also be referred to as an acousticwall effect and results from the interface of the acoustic fieldexerting a strong lateral force (i.e., in the opposite direction to theharvest flow and perpendicular to the axis of the acoustic standingwave) on the suspended particles, thereby keeping the relatively largersized particles from entering the acoustic field and allowing clarifiedfluid (i.e., the fluid containing the smaller-sized product) to enterthe acoustic field, thereby creating an acoustic perfusion cellretention device. In this way, the clarified fluid can escape and thecells are held down by the radiation force. This force is neverpositive, meaning that it always holds the cells down at the interface,i.e., the force is acting in the upstream flow direction, not allowingthe cells to pass through the acoustic interface. The multiple peaks inthe power curve (see discussion of FIG. 36 below) show the existence ofmultiple modes of operation including planar resonance modes andmulti-dimensional modes of operation, indicating that this type ofoperation can be generated through utilization of planar andmulti-dimensional standing waves alike. In systems having 1″×1″dimensions, there exists a planar resonance about every 30 kHz. FIG. 36shows evidence of additional peaks indicating the existence of themulti-dimensional modes. Per unit power, these modes can be equally oreven more effective as the planar resonance modes. As explained above,the cells that are held back by the acoustic radiation force may thenpicked up by the scrubbing motion of the fluid flow field (i.e., therecirculating flow underneath the interface), and be continuouslyreturned to the bioreactor to ensure they receive the nutrition andoxygen to maintain the production of the overall cell culture.

The clarified fluid contains both the desired products and cellfragments, all of which are smaller than whole viable cells. In thisway, the media that is returned to the bioreactor is clarified of cellfragments. Cell fragments absorb media without producing desiredproduct, making the perfusion process less efficient. Thus, there is anefficiency gain and a cost savings obtained by removing these cellfragments using the acoustic perfusion devices of the presentdisclosure. Further clarification of the clarified fluid may be achieveddownstream using a second device or a secondary flow chamber thatcontains another transducer-reflector pair that operates at a differentfrequency. This arrangement traps, clumps, clusters, or agglomeratesparticles having a size of about 10 microns or less that may have passedthrough the original acoustic standing wave, in the same manner asdescribed before. A third transducer-reflector pair operating at anotherfrequency, 3 MHz to 20 MHz, or higher, may be utilized to trap, clump,cluster, or agglomerate and drop out the small cell fragments and debristhat passed through the initial acoustic standing wave and the“interface effect”. This triple-clarified fluid containing the desiredbiomolecules can then directly enter a sterile filter. For example, theoriginal acoustic perfusion device may operate at frequencies up toabout 4 MHz. It is contemplated that the frequency of this second andthird acoustic standing wave field would be from about 6 MHz to about 20MHz, and possibly higher, to trap smaller sized cell fragments.

During startup of a bioreactor at low cell density, e.g., 2 millioncells/mL, the first described mode of operation dominates (FIG. 33, leftimage). As cell density in the bioreactor increases over time, the modeof operation gradually switches from mode 1 to mode 2, and both modesmay coexist at the same time.

When an acoustic standing wave is employed for perfusion in a bioreactorwith an already high cell density, e.g., 50 million cells/mL, the devicetypically starts in the first mode of operation (FIG. 33, left image),until the volume of fluid within the acoustic standing wave isclarified, at which point the operation gradually switches to the seconddescribed mode of operation (FIG. 33, right image). At times, duringoperation, an instability, usually manifested as a perturbation oroscillation of the interface between the two fluids, may growsufficiently strong such that cells enter the volume of fluid within theacoustic standing wave, at which point, for a short period of time, thedevice acts in a combined mode of operation, where both modes are active(i.e., the interface effect prevents cells from entering the acousticfield as explained above, while the acoustic field clarifies the cellsthat have entered the volume of fluid within the acoustic standing wavefield). Once the tightly packed cell clusters have settled out (i.e.,once the volume of fluid within the acoustic standing wave has beensufficiently clarified), the mode of operation is then again that of thesecond described mode of operation, namely, the acoustic interfaceeffect. It is important to note that the device can operate inboth/either of the modes of operation, as described above, withoutexternal switching. In other words, the properties of the fluid streams,e.g., cell concentrations in the streams, and acoustic field dictatewhich mode dominates.

The acoustic standing wave(s) perfusion devices of the presentdisclosure are operated differently compared to prior acoustic filterusages, previously described in literature. Previously, acoustophoresiswas operated such that the protein-producing materials, such as Chinesehamster ovary cells (CHO cells), the most common host for the industrialproduction of recombinant protein therapeutics, were trapped within aplanar ultrasonic standing wave (i.e., remain in a stationary position).Cells were retained in an acoustic field by causing individual cells tomigrate towards the pressure nodal planes of the planar acousticstanding wave. There, as the cells were retained in the standing wave,there was also a physical scrubbing of the cell culture media flowingpast, whereby more cells were trapped as they came in contact with thecells that were already held within the standing wave. The standing waveand harvest fluid flow were then intermittently shut off to allow thecells to drop out of the standing wave and return to the bioreactor.

In contrast, in the present disclosure, the ultrasonic standing wavesare used as a blanket or selector or “force field” in the perfusiondevice. The perfusion device is configured to permit fluid flows thatare against gravity forces. The flows against gravity contribute to thebiological cells sinking. The standing wave is created near the top ofthe filtering device and acts like a filter to prevent the cells fromentering the acoustic field and exiting through the top of the filteringdevice (i.e., acting similar to a force field holding the cells backfrom entering the acoustic field). Thus, two output streams are created,one output stream retaining the cells and exiting through a port at thebottom of the device, and the other output stream being depleted incells and exiting through a port at the top of the device (the cellconcentration in the two output streams being compared to each other).In this mode of operation, there is almost no reliance on clustering,clumping, or agglomeration of the cells within the acoustic field toachieve separation. The combination of fluid dynamics and acoustic fieldto block cells is particularly advantageous in certain applications.Because the cells are not retained in the acoustic field for a period oftime, the filtering device is more easily operated in a continuousarrangement.

Described another way, the acoustic perfusion device has two fluidstreams flowing at different rates. The main fluid stream, carrying theexpressing cell culture, culture media, product, and other bioreactorconstituents, enters the device and is partially diverted into asecondary, lower volume, lower flow fluid stream. This secondary fluidstream passes through the multi-dimensional acoustic standing wave,where the multi-dimensional acoustic standing wave (or generally theinterface effect created by the acoustic standing wave) holds back themain cell culture and allows the expressed biomolecules, the monoclonalantibodies and recombinant proteins, along with other small particlessuch as submicron and micron-sized cell debris, to pass through and befurther collected and processed outside/downstream of the bioreactor.The main fluid stream, containing the main cell culture, is thenrecycled back to the bioreactor. The acoustic standing wave and its“interface effect” can be considered to act as a filter, preventinglarge cells, other particles or bodies, from exiting the bioreactor.

In another application, the acoustic perfusion devices can act as aretention device and cell washing device for cell therapy applications.In continuous cell-culture applications, such as autologous andallogeneic cell therapy, cells may be initially harvested at a very lowcell-density. The cells may be purified, isolated and proliferated inthe cell culture. Relatively few cells may be used to seed a bioreactor.The bioreactor may be used to increase the number of cells in the cellculture. Further processing steps such as concentrating, washing, andmedia exchange are all needed for various applications. The commonalityin all these applications is the need to continuously circulate, add,and/or remove media while retaining cells in a bioreactor (which may betraditional or single-use) with no effect to their viability. Theacoustic cell retention systems described herein operate over a range ofcell recirculation rates, efficiently retain cells over a range ofperfusion (or media removal rates), and can be tuned to fully retain orselectively pass some percentage of cells through fluid flow rate,transducer power or frequency manipulation. Power and flow rates can allbe monitored and used as feedback in an automated control system.Specialty flow paths may also be used such that a small volume of themain fluid flow is “sipped” off and the expressed biomolecules areseparated from the main cell culture.

One advantage of acoustophoresis is that the acoustic radiation forcedoes not harm or negatively affect the biological cells or the desiredbiomolecule product. Moreover, perfusion is continuous, such that thecell culture is kept viable and desired products can be continuallyrecovered therefrom.

In a perfusion bioreactor system, it is desirable to be able to filterand separate the viable biological cells from the expressed materialsthat are in the fluid stream (i.e., cell culture media) and cellulardebris. As previously mentioned, such biological cells may includeChinese hamster ovary (CHO) cells, whose cell genome is manipulated toexpress large biomolecules. Such biomolecules can include recombinantproteins or monoclonal antibodies, and are the desired product to berecovered.

The acoustic perfusion devices of the present disclosure are designed tomaintain a high intensity multi-dimensional acoustic standing wave thatcan act as a filter, permitting smaller particles (such as recombinantproteins or cellular debris) to pass through while excluding largerparticles (such as viable cells). Generally, the device is driven by anoscillator and amplifier (not shown), and the device performance ismonitored and controlled by a computer (not shown). The parameters ofthe energy or signal applied to the transducer, such as, for example,frequency or voltage amplitude, may be controlled and/or modulated. Insome instances it may be desirable to control the applied signal tocompensate for effects such as acoustic streaming. This control may beachieved by amplitude modulation and/or by frequency modulation. Theduty cycle of the propagation of the standing wave may also be utilizedto achieve certain results (i.e. the acoustic beam may be turned on andshut off at different time periods or rates).

FIG. 63 is a log-log graph (logarithmic y-axis, logarithmic x-axis) thatshows the scaling of the radiation forces of cells (MRC-5) andmicrocarriers (Cytodex beads), fluid drag force, and buoyancy force withparticle radius. The graph provides an explanation for the separation ofmicrovesicles, cells or microcarriers from a host fluid based onparameter settings. For example, the frequency, flow rate (drag force)and power can be controlled and/or modulated to prevent or permitpassage of microvesicles, cells or microcarriers. If the fluid flow isincreased from that illustrated in the graph in FIG. 63, the drag forceintersection with the radiation force curves rises along the curves andcan pass the point where particles the size of Cytodex beads arerepelled, so that microvesicles, cells and beads are all passing throughthe acoustic field. At the operating point illustrated in the graph,cells and microvesicles are passed, while the beads are repelled.Lowering the fluid flow until the intersection of the drag force lineintersection with the radiation force curves is below the cell (MRCS)size causes microvesicles (which are smaller than cells and not shown onthe graph) to pass, while cells and beads are repelled or retained bythe acoustic field. Similar effects can be achieved by modulating and/orcontrolling frequency or power of the acoustic field. Thus, theperfusion systems disclosed herein can discriminate based on size with arelatively high level of distinction.

FIG. 1 illustrates a single standing wave system 100 that is comprisedof a reflector plate 101 and an ultrasonic transducer 103 that is set toresonate so as to form a standing wave 102. Excitation frequenciestypically in the range from 100 kHz to 100 MHz are applied by thetransducer 103. One or more multi-dimensional standing waves are createdbetween the transducer 103 and the reflector 101. An ideal standing waveis the sum of two propagating waves that are equal in frequency andintensity and that are traveling in opposite directions, i.e. from thetransducer to the reflector and back. The propagating wavesconstructively interfere with each other and thus generate the standingwave. A dotted line 105 is used to indicate the zero-amplitude of thewave. A node is a point where the wave has minimum amplitude, and isindicated with reference numeral 107. An anti-node is a point where thewave has maximum amplitude, and is indicated with reference numeral 109.The standing wave is illustrated as beginning and ending with localmaxima, however, other implementations are possible. For example, thestanding wave can be offset at the transducer or the reflector so thatlocal minima or maxima are spaced from the transducer or from thereflector. The reflected wave (or wave generated by an opposingtransducer) can be in- or out-of-phase with the transducer generatedwave. The characteristics of the standing wave can be modified and/orcontrolled by the drive signal applied to the transducer, such as bymodifying and/or controlling the phase, amplitude, or frequency of thedrive signal. Acoustically transparent or responsive materials may alsobe used with the transducer or reflector to modify and/or control thestanding wave.

FIG. 2 is a schematic diagram that compares a fed-batch bioreactorsystem 201 (left side) with a perfusion bioreactor system 202 (rightside). Beginning with the fed-batch bioreactor on the left, thebioreactor 210 includes a reaction vessel 220. The cell culture media isfed to the reaction vessel through a feed inlet 222. An agitator 225 isused to circulate the media throughout the cell culture. Here, theagitator is depicted as a set of rotating blades, though any type ofsystem that causes circulation is contemplated. The bioreactor permitsgrowth of a seed culture through a growth/production cycle, during whichtime debris, waste and unusable cells will accumulate in the bioreactorand the desired product (e.g. biomolecules such as monoclonalantibodies, recombinant proteins, hormones, etc.) will be produced aswell. Due to this accumulation, the reaction vessel of a fed-batchprocess is typically much larger than that in a perfusion process. Thedesired product is then harvested at the end of the production cycle.The reaction vessel 220 also includes an outlet 224 for removingmaterial.

Turning now to the perfusion bioreactor 202 on the right-hand side,again, the bioreactor includes a reaction vessel 220 with a feed inlet222 for the cell culture media. An agitator 225 is used to circulate themedia throughout the cell culture. An outlet 224 of the reaction vesselis fluidly connected to the inlet 232 of an acoustic perfusion device230 of the present disclosure, and continuously feeds the bioreactorcontents (containing cells and desired product) to the filtering device.The perfusion device is located downstream of the reaction vessel, andseparates the desired product from the cells. The acoustic perfusiondevice 230 has two separate outlets, a product outlet 234 and a recycleoutlet 236. The product outlet 234 fluidly connects the acousticperfusion device 230 to a containment vessel 240 downstream of theperfusion device, which receives the flow of the desired product (plusmedia) from the perfusion device. From there, furtherprocessing/purification can occur to isolate/recover the desiredproduct. For example, further downstream of this acoustic perfusiondevice may be additional filters such as an ATF, TFF, depth filter,centrifuge, etc. The recycle outlet 236 fluidly connects the acousticperfusion device 230 back to a recycle inlet 226 of the reaction vessel220, and is used to send the cells and cell culture media back into thereaction vessel for continued growth/production. Put another way, thereis a fluid loop between the reaction vessel and the perfusion device.The reaction vessel 220 in the perfusion bioreactor system 202 has acontinuous throughput of product and thus can be made smaller. Thefiltering process is critical to the throughput of the perfusionbioreactor. A poor filtering process implies low throughput and resultsin low yields of the desired product.

FIG. 3 is a cross-sectional view of a generic bioreactor 300 that isuseful for the systems of the present disclosure. As illustrated here,the bioreactor includes a reaction vessel 320 having an internal volume323. A feed inlet 322 at the top of the vessel is used to feed cellculture media into the vessel. An agitator 325 is present. An outlet 324is shown at the bottom of the vessel. A thermal jacket 310 surrounds thereaction vessel, and is used to regulate the temperature of thecells/media. An aerator 312 is located on the bottom of the vessel forproviding gas to the internal volume. Sensors 314 are shown at the topright of the vessel. A pump 316 is illustrated for feeding the cellculture media into the vessel, as is another pump 318 for removing cellculture media from the vessel.

The perfusion systems described above use an acoustic perfusion deviceof the present disclosure. The contents of the bioreactor arecontinuously flowed through the acoustic perfusion device to capture thedesired products.

FIG. 4 is a first embodiment of an acoustic perfusion device 400 thatcan be used with the previously-described systems. The device includesan inlet port 410, an outlet port 430, a first collection port 470, abottom wall 420, and an acoustic chamber 450. The acoustic chamber 450can also be referred to as a fluid cell.

The inlet port 410 is located at a first end 412 of the device.Generally, the inlet port 410 is fluidly connected to an associatedbioreactor and serves as the inlet through which the fluid mixture withcells, fines, and product is introduced to the device. An inlet flowpath 451 leads from the inlet port 410 to the acoustic chamber 450,which contains an internal volume. An upper wall 411 can be presentabove the inlet flow path leading from the inlet port to the acousticchamber, the upper wall having a substantially horizontal orientation.The inlet flow path has a cross-sectional area 452 (illustrated by thedotted square).

The inlet port 410 is located at a first height 402 above the outletport 430, which defines a bottom end of the device. Put another way, theoutlet port 430 is located below the acoustic chamber 450 or below theinlet port 410, or at the bottom end 416 of the device. The placement ofthe outlet port 430 below the inlet port 410 ensures that fluid flowthrough the device is passively urged by gravity towards the outlet port430, and that a hydraulic head is created within the device. The outletport 430 may also be referred to as a fluid recycle port because thehost fluid is recycled or returned from the device to the associatedbioreactor through the outlet port 430. As illustrated here, the outletport 430 is also located at a second end 414 of the device, opposite thefirst end 412. The first end 412 and second end 414 can be considered asbeing opposite ends of an x-axis, while the bottom end 416 and top end418 can be considered as being opposite ends of a z-axis.

The first collection port 470 is located above the acoustic chamber 450at the top end 418 of the device, and is fluidly connected to theacoustic chamber. The device may include additional collection ports,such as second collection port 472, which is spaced apart from the firstcollection port 470. The first and second collection ports 470, 472 aregenerally used to harvest and recover a portion of the desiredbiomolecule byproducts from the device. A collection or harvest flowpath 453 leads from the acoustic chamber to the collection ports 470,472. The collection flow path has a cross-sectional area 454(illustrated by the dotted square). In some particular embodiments, thecross-sectional area 454 of the collection flow path is greater than thecross-sectional area 452 of the inlet flow path. This arrangementprovides one method by which the flow rate of fluid through thecollection ports 470, 472 can be made much lower than the incoming flowrate of fluid. When used in perfusion biomanufacturing, the collectionports can also be referred to as perfusion or harvest ports. Becausefluid depleted in cells and enriched in desired biomolecule products,cell debris, and other fines is harvested, the collection ports can alsobe referred to as harvest ports, and the collection flow path can alsobe referred to as the harvest flow path.

In this embodiment, the bottom wall 420 extends from the inlet port 410to the outlet port 430 of the device. The exact shape of the bottom wall420 can vary to obtain the desired fluid flow. As illustrated here, thebottom wall 420 curves from the inlet port 410 to the outlet port 430 ofthe device. Relative to a line between the inlet port 410 and the outletport 430, illustrated as dotted line 401, the bottom wall 420 has aconcave curve. An outlet flow path 432 leads from the acoustic chamber450 to the outlet port 430.

As illustrated here, a first ultrasonic transducer 460 is located on asidewall 440 of the device at a second height 404 that is above thefirst height 402 (i.e. closer to the top end 418 of the device) andbelow the collection ports 470, 472. This volume above the acousticchamber 450 and below the collection ports 470, 472 is identified hereas a harvest or collection zone 456. The first ultrasonic transducer 460includes a piezoelectric material that can be driven by a drive signalto create a multi-dimensional standing wave in the acoustic chamber 450across the collection flow path 453. An acoustic radiation force fieldthus separates the acoustic chamber 450 from the collection ports 470,472.

In the embodiment of FIG. 4, the device includes a reflector 480 locatedon a wall opposite from the first ultrasonic transducer 460. Thereflector is also located at the second height (i.e. the same height asthe transducer). Together, the transducer 460 and reflector 480 generatea multi-dimensional acoustic standing wave, as illustrated in FIG. 1.

The inlet port 410, outlet port 430, and the collection ports 470, 472are, in this illustrated embodiment, all located on a front wall 475 ofthe device. It is also contemplated that these ports can face in anyother direction, as desired. The front wall 475 is illustrated here ashaving a flat or planar face, and has a constant thickness. However, theshape of the front wall may also vary if desired, for example to changethe cross-sectional areas 452, 454. Finally, the rear wall of the deviceis attached to a mounting piece 490, which contains holes 492 forattaching the perfusion device to a surface for operation.

In use, the fluid mixture containing biological cells and smallermolecules enters the acoustic chamber 450 through the inlet port 410.Inside the acoustic chamber, gravity acts to drag the biological cellsdownwards towards the outlet port 430. A passive settling process occursin the acoustic chamber, creating a fluid with a relatively highconcentration of biological cells at the bottom end 416 of the device,and a fluid with a relatively lower concentration of biological cells atthe top end 418 of the device. The vast majority of incoming fluid, andthereby, the large majority of the cell population never passes throughthe acoustic standing wave(s). The fluid with the high concentration ofbiological cells is pumped back to the bioreactor, and the fluid withthe relatively low concentration of biological cells (and alsocontaining desired biomolecules) is pumped out and collected though thecollection port(s) 470, 472. The acoustic standing wave(s) of the deviceact to prevent significant numbers of biological cells from exitingthrough the collection port(s) 470, 472.

The flow rate through the collection or harvest flow path 453 is, invarious embodiments, at least one order of magnitude smaller than theflow rate through the inlet flow path 451. In more particularembodiments, the flow rate of the fluid mixture entering the devicethrough the inlet port is about 1 liter per minute (L/min) and the flowrate of the fluid depleted in cells exiting the device through thecollection port(s) is about 10 milliliters per minute (mL/min). In sometests, bioreactors having a size of 2 liters to 10 liters have beentested with solutions containing up to 10% yeast and up to 50 millioncells/m L. The flow rate through the inlet port has been from about 0.75L/min to about 3 L/min, with the flow rate through the collection flowpath (i.e. all collection ports together) being about 1 mL/min to about30 mL/min. A 95% cell recovery rate has been achieved.

The acoustic perfusion devices of the present disclosure can filter veryhigh cell densities, around 100 million cells per mL and possibly in therange of about 20 million to about 120 million cells per mL, whereasother filtering technologies such as ATF may be limited to being able tofilter at densities less than 80 million cells per mL. Unlike hollowfiber membranes, the acoustic standing wave(s) can also be tuned toallow passage of cells if desired, as well as allow the passage offines/debris. This tuning for material passage permits the acousticstanding wave(s) to perform a cleaning operation for the bioreactor.Continuous, steady-state operation is possible without pressurefluctuations, and the product stream does not accumulate in thebioreactor or the filtering device.

The acoustic perfusion device can be made of appropriate materials knownin the art. Such materials include high density polyethylene (HDPE),other plastics, and potentially metals and glasses. It has been foundvery convenient for the device to be transparent, so that fluid flow andultrasonic transducer operation can be visually confirmed.

FIG. 5 shows another embodiment of an acoustic perfusion device 500.This embodiment is very similar to the device 400 depicted in FIG. 4.The main difference is that the acoustic perfusion device 500 of FIG. 5has a first ultrasonic transducer 460 on one sidewall of the device anda second ultrasonic transducer 562 on an opposite sidewall 440 thereofin the collection zone 456. Put another way, the two transducers 460,562 are located on opposite sides of the collection flow path 453. Withthis arrangement, the reflector 580 is located within the collectionzone 456 between the first and second ultrasonic transducers 460, 562.The transducers are oriented so that the reflector 580 and first andsecond ultrasonic transducers 460, 562 create multi-dimensional standingwave(s) in the fluid cell 450 as described above, or put another way thetransducers are facing each other. Also illustrated is the outflow pump592 attached to the outlet port 430 of the device, which is used tocontrol the flow rate of the fluid mixture flowing through the device.Not illustrated here is the pump attached to the collection ports (notvisible) of the filtering device 500.

Turning now to FIG. 6, a processing system is shown including anassociated bioreactor 610 and an acoustic perfusion device 630 of thepresent disclosure. The system is set up for use as a perfusionbioreactor. The bioreactor 610 includes a reaction vessel 620 having afeed inlet 622, an outlet 624, and a recycle inlet 626. Fresh media isadded into the feed inlet 622 by an addition pipe 650. Some reactorswill also include an outlet or bleed port (not shown here) to remove or“bleed” cells in order to maintain a constant cell density within areactor. The contents of the reaction vessel (reference numeral 605) aremixed with an agitator 625. The desired product (e.g., recombinantproteins) is continuously produced by cells located within the vessel620, and are present in the media of the bioreactor. The product and thecells in the perfusion bioreactor are drawn from the reaction vesselthrough pipe 652, and enter the acoustic perfusion device 630 throughinlet port 632. Therein, a portion of the desired product is separatedfrom the cells. The desired product can be drawn off through a firstcollection port 634 (which is a product recovery port) and pipe 654 intoa containment vessel 640, or in the case of a truly continuousproduction system, some other downstream purification process. The cellsare returned to the perfusion bioreactor after separation, passing fromoutlet port 636 (which is a fluid recycle port) of the acousticperfusion device through pipe 656 to recycle inlet 626 of the reactionvessel, which form a recycle path. The multi-dimensional standingwave(s) of the acoustic perfusion device are used to create a separationbarrier between the fluid cell of the device and the collection port, sothat a highly reduced number of biological cells are collected incollection port 634.

FIG. 7 and FIG. 8 are views of another exemplary embodiment of anacoustic perfusion device. FIG. 7 is a front cross-sectional view, andFIG. 8 is an exterior perspective view. Notably, this embodiment isspecifically designed such that it can be fabricated with cleanmachining techniques, using Class VI materials (medical device gradeHDPE, for example), or even as single or welded injection molded part.In this manner, this embodiment is an example of a single-use device,which is gamma-stable. The devices are flushed to remove bioburden andthen gamma-irradiated (generally from 25-40 kGy) to sterilize anypotential contamination that could destroy a healthy cell culture, suchas that present in a perfusion bioreactor.

Referring first to FIG. 7, in this device 700, the inlet port 710 andthe collection port 770 are both located at the top end 718 of thedevice, or on the top wall 776 of the device. The outlet port 730 islocated at a bottom end 716 of the device. Here, the inlet port 710 andthe outlet port 730 are both on a first side 712 of the device. Theinlet flow path 751 is in the form of a channel 755 that runs from theinlet port downwards towards the bottom end and past the outlet port,the channel being separated from the acoustic chamber 750 (here, theseparation occurring by an internal wall 756). Fluid will flow downwardsin the channel, then rise upwards into the acoustic chamber 750. Thebottom wall 720 of the acoustic chamber is a sloped planar surface thatslopes down towards the outlet port 730. The location of the ultrasonictransducers 760 are shown here as two squares, between the top end andthe bottom end of the device. The collection flow path 753 is locatedabove the transducers.

Referring now to FIG. 8, the device 700 is shown as being formed withina three-dimensional rectangular housing 706. It can be seen that theoutlet port 730 at the bottom end 716 of the device is located on afront wall 775. Again, the collection port 770 and the inlet port 710are located on a top wall 776. A viewing window 708 made of atransparent material is present in the front wall. Through that viewingwindow, it can be seen that the ultrasonic transducers are mounted inthe rear wall 778 of the device housing. The viewing window acts as areflector to generate the multi-dimensional acoustic standing waves.

FIG. 9 and FIG. 10 are views of yet another exemplary embodiment of anacoustic perfusion device. FIG. 9 is a front cross-sectional view, andFIG. 10 is a perspective view.

Referring first to FIG. 9, in this device 900, there is an inlet port910 present on a front side 975 of the device along the first side 912of the device. An outlet port 930 (best seen in FIG. 10) is locateddirectly opposite and at the same height as the inlet port 910, and isalso located on first side 912. In this embodiment, there is a mainfluid stream that flows almost directly from the inlet port 910 to theoutlet port 930, and the inlet flow path 951 diverts a small side flowinto the acoustic chamber 950 from the first side 912 of the device. Thecollection port 970 is located at the top end 918 of the device, or onthe top wall 976 of the device. A secondary outlet port 980 is locatedon the first side 912 of the device as well, extending from first sidewall 979, and located below the inlet port 910, and can act as a bleedport. The bottom wall 920 of the acoustic chamber is shaped in apyramid-like fashion to taper downwards to a vertex. A drain line 981runs from the bottom of the acoustic chamber 950 to the secondary outletport 980. It is contemplated that here, the secondary outlet port can beused to capture a small flow of highly concentrated cells, which caneither be discarded (cell bleed) or can also be returned back to thebioreactor.

Referring now to FIG. 10, the front wall 975 of the device has arectangular space 960, and the rear wall 978 of the device has arectangular space 962. It is contemplated that one transducer and onereflector can be placed in these two rectangular spaces 960/962 ineither orientation, or that two transducers could be placed in the tworectangular spaces. The inlet port 910 and outlet port 930 are bothvisible in this view. The inlet port 910 is located on the front side ofthe device, and the outlet port 930 is located on the rear side of thedevice (though this arrangement could be reversed if desired). Theclarification flow path 953 is located above the transducers. Althoughnot depicted here, a mounting piece similar to that in FIG. 4 could beattached to the second side 914 of the device.

FIGS. 43-45 are views of yet another exemplary embodiment of an acousticperfusion device. FIG. 43 is a perspective view, FIG. 10 is a pictureshowing a front view, and FIG. 45 is a side view. FIG. 46A/B and FIG.46C/D are schematic front views of possible interior layouts of thedevice. FIG. 46A and FIG. 46B are identical, and are used because of thelarge number of reference numerals. FIG. 46C and FIG. 46D are alsoidentical

Referring now to FIGS. 43-45, in this device 4300, the inlet port 4310and the outlet port 4330 are both located at the bottom end 4316 of thedevice, and the collection port 4370 is located at the top end 4318 ofthe device. The inlet port 4310 is located on a first side 4312 of thedevice, and the outlet port 4330 is located on a second side 4314 of thedevice. In FIG. 43, the outlet port 4330 is attached to a pump 4305,which creates flow through the device 4300. A viewing window 4308 ispresent on the front wall 4375 of the device. The front wall 4375, topwall 4376, rear wall 4378, and first side wall 4379 are part of thehousing 4306 that surrounds the interior of the device.

Referring now to FIG. 43 and FIG. 45, the ultrasonic transducer 4360 islocated on the rear wall 4378 at the top end 4318 of the device. Theviewing window 4308 acts as a reflector to generate themulti-dimensional acoustic standing waves.

In this embodiment, a recirculation pipe 4340 connects the inlet port4310 directly to the outlet port 4330, and forms a recirculation flowpath (arrow 4356) through which cell culture media containing cells andother materials can be continuously recirculated through the perfusiondevice without entering the acoustic chamber 4350. The recirculationpipe 4340 and the recirculation flow path 4356 are located below theacoustic chamber 4350.

An inflow passageway 4380 and an outflow passageway 4390 connect theacoustic chamber 4350 to the recirculation pipe 4340, and split off aportion of the flow of cell culture media from the recirculation pipeinto the acoustic chamber. Arrow 4351 indicates the inlet flow path, andarrow 4355 indicates the outlet flow path. These two passageways areparticularly visible in FIG. 44. Put another way, the inlet flow pathtravels through a different passage than the outlet flow path. Thisarrangement creates a secondary recirculating flow that is tangential tothe acoustic interface, and allows for constant recirculation of cellsbeneath this acoustic interface, traveling in the same net direction asthe recirculation flow path 4356.

The flow geometry of the inflow passageway 4380 and the outflowpassageway 4390 can affect the flow profile through the acousticchamber. FIG. 46A and FIG. 46C are front views showing two differentinternal structures that result in different flow profiles. In these twofigures, the inlet port 4310 is on the right, and the outlet port 4330is on the left.

Considering FIG. 46A and FIG. 46B first, the acoustic chamber 4350 isshown, with the ultrasonic transducer 4360 shown in dashed line. Theacoustic chamber 4350 includes a first side wall 4362 and a second sidewall 4364. The inflow passageway 4380 also has a first wall 4381 and asecond wall 4382, with the first wall 4381 extending beyond the firstside wall 4362, or closer to the inlet port 4310. The bottomcross-sectional area of the inflow passageway (adjacent therecirculation pipe 4340) is indicated by reference numeral 4384, and thetop cross-sectional area of the inflow passageway (adjacent the acousticchamber 4350) is indicated by reference numeral 4383. In embodiments,the top cross-sectional area of the inflow passageway is greater thanthe bottom cross-sectional area of the inflow passageway.

The outflow passageway 4390 also has a first wall 4391 and a second wall4392. The first wall 4391 and the second wall 4392 taper towards eachother from the acoustic chamber 4350 to the recirculation pipe 4340. Thebottom cross-sectional area of the outflow passageway (adjacent therecirculation pipe 4340) is indicated by reference numeral 4394, and thetop cross-sectional area of the outflow passageway (adjacent theacoustic chamber 4350) is indicated by reference numeral 4393. Inembodiments, the top cross-sectional area of the outflow passageway isgreater than the bottom cross-sectional area of the outflow passageway.

It is noted that the top cross-sectional area 4393 of the outflowpassageway is greater than the top cross-sectional area 4383 of theoutflow passageway. The bottom cross-sectional area 4394 of the outflowpassageway is also less than the bottom cross-sectional area 4384 of theoutflow passageway. Desirably, this arrangement promotes the directionfor cells and other larger materials to enter the acoustic chamber 4350,and maximizes their opportunity to exit the acoustic chamber in the samedirection as the main recirculation flow.

Now considering FIG. 46C and FIG. 46D, the first wall 4381 of the inflowpassageway 4380 is essentially in-line with the first side wall 4362.The second wall 4382 is vertical like the first side wall, then widensat the top. The top cross-sectional area 4383 of the inflow passagewayis greater than the bottom cross-sectional area 4384 of the inflowpassageway. The first wall 4391 of the outflow passageway 4390 tapersdownwards, and then becomes vertical. The second wall 4392 tapersinwards from the second side wall 4364 to the recirculation pipe 4340.Again, the top cross-sectional area 4393 of the outflow passageway isgreater than the bottom cross-sectional area 4394 of the outflowpassageway. In FIG. 46C, the top cross-sectional area 4393 of theoutflow passageway is still greater than the top cross-sectional area4383 of the outflow passageway. The bottom cross-sectional area 4394 ofthe outflow passageway can be about equal to or less than the bottomcross-sectional area 4384 of the outflow passageway.

FIGS. 47-49 are views of yet another exemplary embodiment of an acousticperfusion device. FIG. 47 is a perspective view, FIG. 48 is a side view.FIG. 49 is a front schematic view of the interior layout of the device.

Referring now to FIGS. 47-49, in this device 4700, the inlet port 4710and the outlet port 4730 are both located at the bottom end 4716 of thedevice, and the collection port 4770 is located at the top end 4718 ofthe device. The inlet port 4710 is located on the front wall 4775 of thedevice, and the outlet port 4730 is located on the rear wall 4778 of thedevice. In FIG. 47, the outlet port 4730 is attached to a pump 4705,which creates flow through the device 4700. A viewing window 4708 ispresent on the front wall 4775 of the device. The front wall 4775, topwall 4776, and rear wall 4778 are part of the housing 4706 thatsurrounds the interior of the device.

Referring now to FIG. 47 and FIG. 48, the ultrasonic transducer 4760 islocated on the rear wall 4778 at the top end 4718 of the device. Theviewing window 4708 acts as a reflector to generate themulti-dimensional acoustic standing waves.

Again, a recirculation pipe 4740 connects the inlet port 4710 directlyto the outlet port 4730, and forms a recirculation flow path (arrow4756) through which cell culture media containing cells and othermaterials can be continuously recirculated through the perfusion devicewithout entering the acoustic chamber 4750. The recirculation pipe 4740and the recirculation flow path 4756 are located below the acousticchamber 4750.

This embodiment differs from that of FIG. 43 in that a single passageway4772 connects the acoustic chamber 4750 to the recirculation pipe 4740,rather than the two separate passageways (4380, 4390) of FIG. 43.Referring now to FIG. 49, arrow 4751 indicates the inlet flow path, andarrow 4755 indicates the outlet flow path, both traveling through thesingle passageway. This flow still results in a secondary recirculatingflow that is tangential to the acoustic interface, and allows forconstant recirculation of cells beneath this acoustic interface,traveling in the same net direction as the recirculation flow path 4756.

It may be helpful now to describe the ultrasonic transducer(s) used inthe acoustic filtering device in more detail. FIG. 11 is across-sectional diagram of a conventional ultrasonic transducer. Thistransducer has a wear plate 50 at a bottom end, epoxy layer 52, ceramicpiezoelectric element 54 (made of, e.g. Lead Zirconate Titanate (PZT)),an epoxy layer 56, and a backing layer 58. On either side of the ceramicpiezoelectric element, there is an electrode: a positive electrode 61and a negative electrode 63. The epoxy layer 56 attaches backing layer58 to the piezoelectric element 54. The entire assembly is contained ina housing 60 which may be made out of, for example, aluminum. Thehousing is used as the ground electrode. An electrical adapter 62provides connection for wires to pass through the housing and connect toleads (not shown) which attach to the piezoelectric element 54.Typically, backing layers are designed to add damping and to create abroadband transducer with uniform displacement across a wide range offrequency and are designed to suppress excitation of particularvibrational eigen-modes of the piezoelectric element. Wear plates areusually designed as impedance transformers to better match thecharacteristic impedance of the medium into which the transducerradiates.

FIG. 12 is a cross-sectional view of an ultrasonic transducer 81according to an example of the present disclosure, which is used in theacoustic filtering device of the present disclosure. Transducer 81 isshaped as a disc or a plate, and has an aluminum housing 82. Thealuminum housing has a top end and a bottom end. The transducer housingmay also be composed of plastics, such as medical grade HDPE or othermetals. The piezoelectric element is a mass of perovskite ceramic, eachconsisting of a small, tetravalent metal ion, usually titanium orzirconium, in a lattice of larger, divalent metal ions, usually lead orbarium, and O²⁻ ions. As an example, a PZT (lead zirconate titanate)piezoelectric element 86 defines the bottom end of the transducer, andis exposed from the exterior of the bottom end of the housing. Thepiezoelectric element is supported on its perimeter by a small elasticlayer 98, e.g. epoxy, silicone or similar material, located between thepiezoelectric element and the housing. Put another way, no wear plate orbacking material is present. However, in some embodiments, there is alayer of plastic or other material separating the piezoelectric elementfrom the fluid in which the acoustic standing wave is being generated.The piezoelectric element/crystal has an exterior surface (which isexposed) and an interior surface as well. In particular embodiments, thepiezoelectric element/crystal is an irregular polygon, and in furtherembodiments is an asymmetrical irregular polygon.

Screws 88 attach an aluminum top plate 82 a of the housing to the body82 b of the housing via threads. The top plate includes a connector 84for powering the transducer. The top surface of the PZT piezoelectricelement 86 is connected to a positive electrode 90 and a negativeelectrode 92, which are separated by an insulating material 94. Theelectrodes can be made from any conductive material, such as silver ornickel. Electrical power is provided to the PZT piezoelectric element 86through the electrodes on the piezoelectric element. Note that thepiezoelectric element 86 has no backing layer or epoxy layer. Putanother way, there is an interior volume or an air gap 87 in thetransducer between aluminum top plate 82 a and the piezoelectric element86 (i.e. the housing is empty). A minimal backing 58 (on the interiorsurface) and/or wear plate 50 (on the exterior surface) may be providedin some embodiments, as seen in FIG. 13.

The transducer design can affect performance of the system. A typicaltransducer is a layered structure with the ceramic piezoelectric elementbonded to a backing layer and a wear plate. Because the transducer isloaded with the high mechanical impedance presented by the standingwave, the traditional design guidelines for wear plates, e.g., halfwavelength thickness for standing wave applications or quarterwavelength thickness for radiation applications, and manufacturingmethods may not be appropriate. Rather, in one embodiment of the presentdisclosure the transducers, there is no wear plate or backing, allowingthe piezoelectric element to vibrate in one of its eigenmodes with ahigh Q-factor, or in a combination of several eigenmodes. The vibratingceramic piezoelectric element/disk is directly exposed to the fluidflowing through the fluid cell.

Removing the backing (e.g. making the piezoelectric element air backed)also permits the ceramic piezoelectric element to vibrate at higherorder modes of vibration with little damping (e.g. higher order modaldisplacement). In a transducer having a piezoelectric element with abacking, the piezoelectric element vibrates with a more uniformdisplacement, like a piston. Removing the backing allows thepiezoelectric element to vibrate in a non-uniform displacement mode. Thehigher order the mode shape of the piezoelectric element, the more nodallines the piezoelectric element has. The higher order modal displacementof the piezoelectric element creates more trapping lines, although thecorrelation of trapping line to node is not necessarily one to one, anddriving the piezoelectric element at a higher frequency will notnecessarily produce more trapping lines.

It is contemplated that, in some embodiments of the acoustic filteringdevice of the present disclosure, the piezoelectric element may have abacking that minimally affects the Q-factor of the piezoelectric element(e.g. less than 5%). The backing may be made of a substantiallyacoustically transparent material such as balsa wood, foam, or corkwhich allows the piezoelectric element to vibrate in a higher order modeshape and maintains a high Q-factor while still providing somemechanical support for the piezoelectric element. The backing layer maybe a solid, or may be a lattice having holes through the layer, suchthat the lattice follows the nodes of the vibrating piezoelectricelement in a particular higher order vibration mode, providing supportat node locations while allowing the rest of the piezoelectric elementto vibrate freely. The goal of the lattice work or acousticallytransparent material is to provide support without lowering the Q-factorof the piezoelectric element or interfering with the excitation of aparticular mode shape.

Placing the piezoelectric element in direct contact with the fluid alsocontributes to the high Q-factor by avoiding the dampening and energyabsorption effects of the epoxy layer and the wear plate. Otherembodiments of the transducer(s) may have wear plates or a wear surfaceto prevent the PZT, which contains lead, from contacting the host fluid.This additional layer may be desirable in, for example, biologicalapplications such as separating blood, biopharmaceutical perfusion, orfed-batch filtration of mammalian cells. Such applications might use awear layer such as chrome, electrolytic nickel, or electroless nickel.Chemical vapor deposition could also be used to apply a layer ofpoly(p-xylylene) (e.g. Parylene) or other polymer. Organic andbiocompatible coatings such as silicone or polyurethane are also usableas a wear surface. Thin films, such as a PEEK film, can also be used asa cover of the transducer surface exposed to the fluid with theadvantage of being a biocompatible material. In one embodiment, the PEEKfilm is adhered to the face of the piezomaterial using pressuresensitive adhesive (PSA). Other films can be used as well.

In some embodiments, for applications such as oil/water emulsionsplitting and others such as perfusion, the ultrasonic transducer has anominal 2 MHz resonance frequency. Each transducer can consume about 28W of power for droplet trapping at a flow rate of 3 GPM. This powerconsumption translates to an energy cost of 0.25 kW hr/m³. This cost isan indication of the very low cost of energy of this technology.Desirably, each transducer is powered and controlled by its ownamplifier. In other embodiments, the ultrasonic transducer uses a squarepiezoelectric element, for example with 1″×1″ dimensions. Alternatively,the ultrasonic transducer can use a rectangular piezoelectric element,for example with 1″×2.5″ dimensions. Power dissipation per transducerwas 10 W per 1″×1″ transducer cross-sectional area and per inch ofacoustic standing wave span in order to get sufficient acoustic trappingforces. For a 4″ span of an intermediate scale system, each 1″×1″ squaretransducer consumes 40 W. The larger 1″×2.5″ rectangular transducer uses100 W in an intermediate scale system. The array of three 1″×1″ squaretransducers would consume a total of 120 W and the array of two 1″×2.5″transducers would consume about 200 W. Arrays of closely spacedtransducers represent alternate potential embodiments of the technology.Transducer size, shape, number, and location can be varied as desired togenerate desired multi-dimensional acoustic standing wave patterns.

In some examples, the size, shape, and thickness of the transducer candetermine the transducer displacement at different frequencies ofexcitation. Transducer displacement with different frequencies mayaffect separation efficiency. Typically, the transducer is operated atfrequencies near the thickness resonance frequency (half wavelength).Gradients in transducer displacement typically result in more trappinglocations for the cells/biomolecules. Higher order modal displacementscan generate three-dimensional acoustic standing waves with stronggradients in the acoustic field in all directions, thereby creatingstrong acoustic radiation forces in all directions, which forces may,for example, be equal in magnitude, leading to multiple trapping lines,where the number of trapping lines correlate with the particular modeshape of the transducer.

To investigate the effect of the transducer displacement profile onacoustic trapping force and separation efficiencies, an experiment wasrepeated ten times using a 1″×1″ square transducer, with all conditionsidentical except for the excitation frequency. Ten consecutive acousticresonance frequencies, indicated by circled numbers 1-9 and letter A onFIG. 14, were used as excitation frequencies. The conditions wereexperiment duration of 30 min, a 1000 ppm oil concentration ofapproximately 5-micron SAE-30 oil droplets, a flow rate of 500 ml/min,and an applied power of 20 W. Oil droplets were used because oil is lessdense than water, and can be separated from water using acoustophoresis.

FIG. 14 shows the measured electrical impedance amplitude of a squaretransducer as a function of frequency in the vicinity of the 2.2 MHztransducer resonance. The minima in the transducer electrical impedancecorrespond to acoustic resonances of the water column and representpotential frequencies for operation. Additional resonances exist atother frequencies where multi-dimensional standing waves are excited.Numerical modeling has indicated that the transducer displacementprofile varies significantly at these acoustic resonance frequencies,and thereby directly affects the acoustic standing wave and resultingtrapping force. Since the transducer operates near its thicknessresonance, the displacements of the electrode surfaces are essentiallyout of phase. The typical displacement of the transducer electrodes maynot be uniform and varies depending on frequency of excitation. As anexample, at one frequency of excitation with a single line of trappedoil droplets, the displacement has a single maximum in the middle of theelectrode and minima near the transducer edges. At another excitationfrequency, the transducer profile has multiple maxima leading tomultiple trapped lines of oil droplets. Higher order transducerdisplacement patterns result in higher trapping forces and multiplestable trapping lines for the captured oil droplets.

As the oil-water emulsion passed by the transducer, the trapping linesof oil droplets were observed and characterized. The characterizationinvolved the observation and pattern of the number of trapping linesacross the fluid channel, as shown in FIG. 15, for seven of the tenresonance frequencies identified in FIG. 14. Different displacementprofiles of the transducer can produce different (more) trapping linesin the standing waves, with more gradients in displacement profilegenerally creating higher trapping forces and more trapping lines.Different displacement profiles of the transducer can produce different(more) trapping lines in the standing waves, with more gradients indisplacement profile generally creating higher trapping forces and moretrapping lines. It is noted that although the different trapping lineprofiles shown in FIG. 15 were obtained at the frequencies shown in FIG.14, these trapping line profiles can also be obtained at differentfrequencies. FIG. 15 shows the different crystal vibration modespossible by driving the crystal to vibrate at different fundamentalfrequencies of vibration. The 3D mode of vibration of the crystal iscarried by the acoustic standing wave across the fluid in the chamberall the way to the reflector and back. The resulting multi-dimensionalstanding wave can be thought of as containing two components. The firstcomponent is a planar out-of-plane motion component (uniformdisplacement across crystal surface) of the crystal that generates astanding wave, and the second component is a displacement amplitudevariation with peaks and valleys occurring in lateral directions acrossthe crystal surface. Three-dimensional force gradients are generated bythe standing wave. These three-dimensional force gradients result inlateral radiation forces that stop and trap the particles with respectto the flow by overcoming the viscous drag force. In addition, thelateral radiation forces are responsible for creating tightly packedclusters of particles. Therefore, particle separation and gravity-drivencollection depends on generating a multi-dimensional standing wave thatcan overcome the particle drag force as the mixture flows through theacoustic standing wave. Multiple particle clusters are formed alongtrapping lines in the axial direction of the standing wave, as presentedschematically in FIG. 15.

FIG. 16 is a numerical model showing a pressure field that matches the 9trapping line pattern. The numerical model is a two-dimensional model;and therefore three trapping lines are observed. Two more sets of threetrapping lines exist in the third dimension perpendicular to the planeof the page.

The lateral force of the acoustic radiation force generated by thetransducer can be increased by driving the transducer in higher ordermode shapes, as opposed to a form of vibration where the crystaleffectively moves as a piston having a uniform displacement. Theacoustic pressure is proportional to the driving voltage of thetransducer. The electrical power is proportional to the square of thevoltage. The transducer is typically a thin piezoelectric plate, withelectric field in the z-axis and primary displacement in the z-axis. Thetransducer is typically coupled on one side by air (i.e., the air gapwithin the transducer) and on the other side by the fluid mixture of thecell culture media. The types of waves generated in the plate are knownas composite waves. A subset of composite waves in the piezoelectricplate is similar to leaky symmetric (also referred to as compressionalor extensional) Lamb waves. The piezoelectric nature of the platetypically results in the excitation of symmetric Lamb waves. The wavesare leaky because they radiate into the water layer, which result in thegeneration of the acoustic standing waves in the water layer. Lamb wavesexist in thin plates of infinite extent with stress free conditions onits surfaces. Because the transducers of this embodiment are finite innature, the actual modal displacements are more complicated.

FIG. 17 shows the typical variation of the in-plane displacement(x-displacement) and out-of-plane displacement (y-displacement) acrossthe thickness of the plate, the in-plane displacement being an evenfunction across the thickness of the plate and the out-of-planedisplacement being an odd function. Because of the finite size of theplate, the displacement components vary across the width and length ofthe plate. In general, a (m,n) mode is a displacement mode of thetransducer in which there are m undulations in transducer displacementin the width direction and n undulations in the length direction, andwith the thickness variation as described in FIG. 17. The maximum numberof m and n is a function of the dimension of the crystal and thefrequency of excitation. Additional three-dimensional modes exist thatare not of the form (m,n).

The transducers are driven so that the piezoelectric element vibrates inhigher order modes of the general formula (m, n), where m and n areindependently 1 or greater. Generally, the transducers will vibrate inhigher order modes than (2,2). Higher order modes will produce morenodes and antinodes, result in three-dimensional standing waves in thewater layer, characterized by strong gradients in the acoustic field inall directions, not only in the direction of the standing waves, butalso in the lateral directions. As a consequence, the acoustic gradientsresult in stronger trapping forces in the lateral direction.

The transducer can be driven by a signal, such as a voltage signal, acurrent signal, a magnetic signal, an electromagnetic signal, acapacitive signal, or any other type of signal to which the transduceris responsive to create a multi-dimensional acoustic standing wave. Inembodiments, the voltage signal driving the transducer can have asinusoidal, square, sawtooth, pulsed, or triangle waveform; and have afrequency of 500 kHz to 10 MHz. The voltage signal can be driven withpulse width modulation, which produces any desired waveform. The voltagesignal can also have amplitude or frequency modulation start/stopcapability to eliminate streaming.

The transducers are used to create a pressure field that generatesacoustic radiation forces of the same order of magnitude both orthogonalto the standing wave direction and in the standing wave direction. Whenthe forces are roughly the same order of magnitude, particles of size0.1 microns to 300 microns will be moved more effectively towards“trapping lines”, so that the particles will not pass through thepressure field and continue to exit through the collection ports of thefiltering device. Instead, the particles will remain within the acousticchamber to be recycled back to the bioreactor.

In biological applications, it is contemplated that all of the parts ofthe system (i.e., the bioreactor, acoustic filtering device, tubingfluidly connecting the same, etc.) can be separated from each other andbe disposable. Avoiding centrifuges and filters allows better separationof the CHO cells without lowering the viability of the cells. Thetransducers may also be driven to create rapid pressure changes toprevent or clear blockages due to agglomeration of CHO cells. Thefrequency of the transducers may also be varied to obtain optimaleffectiveness for a given power.

The following examples are provided to illustrate the devices,components, and methods of the present disclosure. The examples aremerely illustrative and are not intended to limit the disclosure to thematerials, conditions, or process parameters set forth therein.

EXAMPLES Example 1

FIG. 18 shows an experimental setup for an acoustic perfusion device asdescribed in detail above. This acoustic perfusion device is verysimilar to that illustrated in FIG. 5, except the bottom wall is notcurved, but rather runs horizontally from the first end and then anglesdirectly to the outlet port. Tubes are connected to the inlet port,outlet port, and the two collection ports. A pump is also visiblyfluidly connected to the outlet port.

FIG. 19 is a picture of another acoustic perfusion device of the presentdisclosure, similar to the embodiment shown in FIG. 5, having twoultrasonic transducers and a concave bottom wall leading from the inletport to the outlet port at the bottom end of the device. A cellcontaining fluid mixture is also present in the device. In this picture,acoustic standing waves are created in the collection zone between thereflector and first and second transducers as described above. Theacoustic field generated thereby is indicated by waves and referencenumeral 1664. The flow pattern of the fluid mixture through the devicefrom the inlet port to the outlet port is shown with an arrow (referencenumeral 1610) indicating the direction of fluid flow into the device andarrows (reference numeral 1630) indicating the direction of fluid flowthrough the device towards the outlet port. Finally, the general flowpattern of the desired product out of the device through the first andsecond collection ports is shown with arrows (reference numeral 1670)indicating the direction of flow.

Acoustophoretic separation has been tested using the acoustic perfusiondevice of FIG. 19 and the methods of separation of the presentdisclosure on different lines of Chinese hamster ovary (CHO) cells.FIGS. 20-28 show various test results varying different parameters andmeasuring different values using a Beckman Coulter Cell ViabilityAnalyzer.

The perfusion flow rates with the acoustic filtering device were fromabout 2 mL/min to about 10 mL/min, or the flow rates were about 1 VVD toabout 5 VVD for a 2.7 L working volume bioreactor. The VVD refers to the“vessel volume per day”, or how many times the volume of the bioreactorvessel is cycled through the acoustic filtering device in one day. Theperfusion flow rate (Qp) was collected through the perfusion ports. Incontrast, the feed flow rates (Qf) were from about 40 mL/min to about200 m L/m in.

The feed solution had a starting CHO cell density of 50×10⁶ cells/mL.The reactor size was 2.7 L and the feed volume of the host fluid was 1.5L. In total, a series of seven tests (T1-T7) were performed to study theeffect of varying the VVD and flow split in a 2.7 L volume reactor. Theparameters for the tests are shown in Table 1 below.

TABLE 1 System results for a 2.7 L reactor and feed volume of 1.5 L FlowSplit Qp (Qp/Qf) 1 VVD 1.5 VVD 2 VVD 5.2 VVD Qf 5.0% T1 T2 T3 T7 2.5% T4T5 T6

The results included a cell clarification efficiency between 89-93% at aDC voltage of 45V, regardless of the flow rate as shown in FIG. 20. Itis noted that the DC voltage for T1 was fixed at 60V, whereas for testsT2-T7 the DC voltage was reduced to a fixed amount of 45V. Thetransducer voltage amplitude is estimated to be about half of thesevalues.

The results further included a perfusate turbidity reduction of 90-94%compared to the feed, as shown in FIG. 21. This figure shows theturbidity of the feed, the recirculated fluid (Qc), and the perfusate(Qp). The feed entered the inlet port, the recirculated fluid exited theoutlet port and was recirculated, and the perfusate exited the perfusionport of the acoustic filtering device. It is noted that the turbiditymeasurements for tests T1 and T2 resulted in a hardware error, and theirresults are not displayed. The results of tests T3-T7 are displayed,which equated to a 6-10% turbidity in the perfusion stream relative tothe feed stream, regardless of flow rate.

FIG. 22 was produced by a Beckman Coulter Cell Viability Analyzer andrevealed a cell viability for each flow rate that was within the errorof instrument (i.e., ±6%), with the control ranging from 79-84% over alltests.

Further testing was performed using a solution designated “CHO Line A”.The solution had a starting cell density of 50×10⁶ cells/mL, a turbidityof 2,400 NTU, and cell viability of roughly 80%. The solution wasseparated using a device of the present disclosure in a system having areactor size of 2.7 L. The volume of the feed fluid was between 1.5 Land 2.0 L. The perfused flow rates were from 2 mL/min to 10 mL/min, orfrom 1 to 5 VVD. A series of six tests were performed to study theeffect of varying the VVD and flow split on acoustic filtrationperformance for the 2.7 L volume reactor. The parameters for the testsare shown in Table 2 below.

TABLE 2 System results for a 2.7 L reactor and feed volume from 1.5L-2.0 L T1 T2 T3 VVD 1.5 VVD 2 VVD 1 Flow Split 5.00% Flow Split 5.00%Flow Split 2.50% Perfused 2.8 Perfused 3.8 Perfused 1.9 Flow Flow Flow(ml/min) (ml/min) (ml/min) Feed Flow 56 Feed Flow 75 Feed Flow 75(ml/min) (ml/min) (ml/min) T4 T5 T6 VVD 1.5 VVD 2 VVD 5.2 Flow Split2.50% Flow Split 2.50% Flow Split 5.00% Perfused 2.8 Perfused 3.8Perfused 10 Flow Flow Flow (ml/min) (ml/min) (ml/min) Feed Flow 112.5Feed Flow 150 Feed Flow 194.2 (ml/min) (ml/min) (ml/min)

FIG. 23 shows the measured total cell density of feed flow,recirculation flow, and perfused flow. The bioreactor cell retention forthe tests shows an approximately 90% perfusion efficiency. FIG. 24 showsthe measured cell viability for the tests, revealing no significantchange in viability across the tests.

Next, additional testing was performed using a solution designated “CHOLine B”. The solution had a starting cell density of 75×10⁶ cells/mL, aturbidity of 2,300 NTU, and cell viability of roughly 80%. The solutionwas separated using a device of the present disclosure in a systemhaving a reactor size of 2.7 L. Four tests were performed (T1-T4). Twoof the tests (T1, T3) used a device having a single transducer. Theother two tests (T2, T4) used a device having two transducers in series(such that the fluid ran through both standing waves). The parametersfor the tests are shown in Table 3 below.

TABLE 3 System results for a 2.7 L reactor and feed volume from 1.5L-2.0 L T1 T2 Transducers 1 Transducers 2 VVD 1 VVD 1 Perfused Flow(mL/min) 1.9 Perfused Flow (mL/min) 1.9 Feed Flow (mL/min) 75 Feed Flow(mL/min) 75 T3 T4 Transducers 1 Transducers 2 VVD 2 VVD 2 Perfused Flow(mL/min) 3.8 Perfused Flow (mL/min) 3.8 Feed Flow (mL/min) 150 Feed Flow(mL/min) 150

FIG. 25 shows the measured total cell density of feed, recirculation,and perfusion flows. The bioreactor cell retention for the tests shows aperfusion efficiency greater than 90%. The results further evidenced anapproximately 3-5% greater efficiency when using two transducers ratherthan a single transducer. FIG. 26 shows the measured cell viability forthe tests, revealing no significant change in viability across thetests. Practically speaking, operating at low VVD offers a number ofadvantages, such as media cost reduction.

Example 2

FIG. 27 shows another experimental setup for an acoustic perfusiondevice similar to that illustrated in FIG. 8. Tubes are connected to theinlet port, outlet port, and the collection port.

The device was tested at a transducer voltage of 40V peak to peak, aperfused flow rate of 15-30 mL/min, and a recirculation flow rate of 2L/min. Samples were taken every 45-60 minutes, and the cell retentionrate was determined. FIG. 28 shows the results. The y-axis is theretention, stated in terms of percentage (calculated by comparing theoutput cell count with the input, or bioreactor/culture cell count). Thex-axis is both the applied DC voltage (in V) and the perfused, orharvest, flow rate (in mL/min); it is merely coincidental that the rangeof numeric values for V and mL/min are the same. The cell retentionefficiency remained above 95% for perfused flow rates up to 20 mL/min,and remained above 90% up to about 25 mL/min. FIG. 33 is a compositepicture showing the device in a startup or cell settling mode (left) andin a steady state cell retention mode (right).

Next, experiments were performed to determine what factors would affectcell retention. The perfused flow rate was varied, as was the transducervoltage. When the perfused flow rate was varied, the transducer voltagewas maintained at 40V peak to peak and the recirculation flow rate wasmaintained at 2 L/min. When the transducer voltage was varied, theperfused flow rate was maintained at 20 mL/min and the recirculationflow rate was maintained at 2 L/min. The results indicated that, forthis particular embodiment, a perfused flow rate of about 15 mL/min toabout 28 mL/min was optimum, and a transducer voltage of about 15V peakto peak to about 28V peak to peak was optimum.

A better understanding of the added functionality provided by anacoustic perfusion device can be demonstrated by examining the observedcell samples coming in and being harvested from the device. FIG. 29A isa microscope image (from a Vi-Cell cell counter) of the feed suspension,here a viable cell culture population with approximately 56 millioncells/mL. Several large, round, healthy cells can be observed. FIG. 29Bis a histogram showing the distribution of cell diameter in thepopulation. The diameter distribution is strongly bi-modal, aroundvalues of about 11 microns and about 23 microns. These two modescorrespond roughly with the smaller debris and non-viable cells, and thelarger viable cells. It should be noted that this sample is from aparticularly “dirty” cell population. In general, a production cell linewould be far cleaner, and the peak at approximately 11 microns would bemuch smaller, or even non-existent.

FIG. 30A is another microscope image (from a Vi-Cell cell counter), thistime of the flow harvested from the acoustic perfusion device. In thisimage, very few bright, large cells are observed, in contrast to FIG.29A. Rather, the image is filled with more smaller, darker particles, ordebris. The experimental conditions in this case were a perfused rate of4 mL/min and a recirculation rate of 2 L/min with a DC input voltage of30V. This qualitative observation is confirmed by the histogram in FIG.30B, which shows the distribution of diameters in the perfusate. Lookingat FIG. 30B, the particle distribution is now unimodal, with the peak ofapproximately 9 microns. This distribution indicates that the larger,viable cells have been trapped and retained, or otherwise largelyprevented from exiting in the perfusate. Smaller cells are passingthrough, together with cell debris, fines, and fragments, from submicronto micron sized.

A computational fluid dynamics (CFD) model was made of this device. FIG.31 shows the velocity distributions within the device after 500 seconds.The units are in meters/second (m/s). As expected, the highestvelocities are found in the channel leading downwards from the inletport to the outlet port. The velocity is near zero in the fluid cell andout through the collection port. This velocity profile is important fortwo reasons: the acoustic field is more effective in a flow with alower, more uniform velocity, and because the cells used inbio-manufacturing are sensitive to flow, and the induced shear rate.

FIG. 32 is a diagram illustrating several aspects of this embodiment.Fluid flows into the device through the inlet port 710 (arrow 780) andinto the acoustic chamber 790. The volume of fluid 750 below theacoustic chamber contains the tangential flow path, indicated by arrow782. Fluid with a relatively high amount of viable cells will exitthrough the outlet port 730, as indicated by arrow 781. The acousticinterface effect/region created by the standing waves is marked withreference numeral 783, and is upstream of the acoustic standing wavefield 784. The acoustic interface roughly coincides with an x-y plane inthis diagram. This interface effect separates large cells from smallercell fragments, particulate debris, desired biomolecules, etc., whichcan pass through the interface 783 and the acoustic standing wave field784. By way of comparison, the cell aggregates that arise within theacoustic standing wave field 784 during the first mode of operation (seeFIG. 41) can be described as being aligned in the y-z plane. Inoperation, the separation caused by the interface effect occurs at theinterface region 783 as any large cells are held back by the “interface”or “barrier” effect. The harvest flow stream 785 containing the smallerfragments, particulate debris, desired biomolecules, etc. then exitsthrough harvest port 770. The tangential flow path is part of the inletflow path, and is located below the interface region 783 generated bythe acoustic standing wave. The tangential flow path will transport awayboth the clusters of cells that drop from the acoustic standing wavefield 784 due to gravity effects and the cells that are retained by theacoustic interface effect.

FIG. 61 shows a device 6100 that is substantially similar to the device700 of FIG. 7 and FIG. 32. The same numbering is used in FIG. 61. Themain difference between the device illustrated in FIG. 32 and thatillustrated in FIG. 61 is the flow of particles near transducers 760,which is illustrated with arrow 6153, which runs down past thetransducers 760 and down to acoustic interface effect/region 783. Thisflow pattern near transducers 760 occurs because the particles in FIG.61 are too small or have an acoustic contrast factor that isinsufficient to be affected by the acoustic standing wave created bytransducers 760 (and the reflector), and so they can reenter theacoustic standing wave. This behavior may also be affected by thebuoyancy characteristics of the particles and the fluid flow of thesystem, and thus the Stokes drag experienced by the particles.

Example 3

Another way of explaining the operation of the acoustic perfusion devicecan be understood by looking at the results of a numerical study. In thenumerical study, two fluids with differing effective acoustic properties(i.e., speed of sound and density), were modeled with an interfacebetween them in COMSOL, a numerical simulation software. The acousticfield is calculated and therefrom the lateral radiation force acting ona particle in the direction of the fluid velocity is calculated usingGorkov's equation.

FIG. 34 shows the geometry of the simulation, utilizing a piezoelectrictransducer, steel reflector, aluminum housing, and two fluids: the firstfluid being water within the acoustic field, simulating the clarifiedfluid, and the second fluid being a 15% concentration of CHO cells inwater solution outside of the acoustic field, the second fluid having ahigher density and higher speed of sound than the water fluid andsimulating the bioreactor fluid containing the cells.

The two fluids were separated as indicated by the solid line in themodel of FIG. 34. In this setup, the fluid velocity through the systemwas in a horizontal direction from left to right. Therefore, in order toact as a retention device, the acoustic field needs to generate a forceon the cells that acts in the negative x-direction (i.e, opposite thefluid velocity). Water was modeled with a fluid density of 1000 kg/m³and a speed of sound of 1500 m/s. CHO cells were modeled having adensity of 1050 kg/m³ and a speed of sound of 1550 m/s. A coupledmulti-physics numerical simulation that included a full piezoelectricsimulation of the piezoelectric material, an acoustic simulation of thetwo fluids, and a linear elastic simulation in the steel and aluminumbodies was performed at various frequencies of excitation. Thetransducer was driven at a peak voltage of 40 V.

FIGS. 35A-35C show the acoustic pressure in the two fluids and thedisplacement of the piezoelectric material, the aluminum housing, andthe steel reflector of the model at frequencies of operation of 2.218MHz, 2.2465 MHz, and 2.3055 MHz. The lateral radiation force (i.e.,horizontally in the direction of the fluid flow), was calculated at theinterface between the two fluids along with real electrical powerconsumed by the transducer. The structural displacement of thetransducer and steel are shown, along with the acoustic pressure in thefluid.

FIG. 36 shows the lateral radiation force (N) and the lateral radiationforce normalized by power (N/W) versus frequency acting on the suspendedCHO cells. This graph shows that at the resonance frequencies (i.e.,local maxima in power), the average lateral radiation force on theinterface is negative, meaning that it is in the negative-x direction.The result is the creation of an acoustic barrier effect or an acousticinterface effect. That is, the acoustic field at the interface betweenthe two fluids exerts a strong lateral force on the suspended particlesin a direction opposite to the fluid flow, thereby keeping the largerparticles from entering the acoustic field and allowing the first fluid(i.e., fluid containing smaller particles, such as the desired product,and excluding whole cells) to enter the acoustic field, thereby creatingan acoustic perfusion cell retention device. In this way, the clarifiedfluid can escape and the cells are held back by the radiation force.This force is never positive, meaning that it always holds the cellsback at the interface, not allowing them to cross the interface. Themultiple peaks in the power curve show the existence of multiple modesof operation including planar resonance modes and multi-dimensionalmodes of operation, indicating that this type of operation can begenerated through utilization of planar and multi-dimensional standingwaves alike. In systems having 1″×1″ dimensions, there exists a planarresonance about every 30 kHz. The graph shows evidence of additionalpeaks indicating the existence of the multi-dimensional modes. Per unitpower, these modes can be equally or even more effective as the planarresonance modes. As explained above, the cells that are held back by theacoustic radiation force may then be picked up by the scrubbing motionof the flow field (i.e., the recirculating flow underneath theinterface), and be continuously returned to the bioreactor to ensurethey receive the nutrition and oxygen to maintain the production of theoverall cell culture.

Example 4

FIG. 37 and FIG. 38 show another experimental setup for an acousticperfusion device similar to that illustrated in FIG. 9. Tubes areconnected to the inlet port, outlet port, the collection port, and thesecondary outlet port (for a flow concentrated cells). Arrows areincluded to illustrate fluid flow. Arrows indicate the flow into theinlet port; the flow out of the outlet port; the perfusate flow out thetop of the device and the flow of concentrate out the bottom of thedevice. The flow through the inlet port to the outlet port is therecirculation flowrate. The perfusate flow out the top of the device isthe perfused flowrate containing clarified fluid depleted in cells andcontaining desired product. The flow of concentrate out the bottom ofthe device is the concentrated cell flow. The concentrated cell flow canbe used for a cell bleed operation or if desired, the cells can bereturned to the bioreactor.

The device was tested at a transducer voltage of 40V peak to peak, aperfused flow rate (out the top) of 1-10 mL/min, a recirculation flowrate of 0.75-1 L/min, and a concentrate flow rate (out the bottom) of 15mL/min. The cell retention rate was determined for different perfusedflowrates. FIG. 39 shows the results. The y-axis is the retention with1.00 indicating 100% retention. The cell retention efficiency remainedabove 98% for perfused flow rates up to 7 mL/min, and was just below 90%at 10 mL/min.

Example 5

The device of FIG. 7 and FIG. 8 was tested at different flowrates on twodifferent days. FIG. 50 shows the cell retention (%) versus time. Here,higher values are more desirable, and most values were over 95% atflowrates ranging from 1.5 mL/min to 8 mL/min. FIG. 51 shows theperfusate cell density (million cells/mL) versus time. Here, lowervalues are more desirable (indicating successful cell separation). Asexpected, results were better at lower flowrates.

Example 6

A computational fluid dynamics (CFD) model was made of the device withthe internal structure of FIG. 46A. FIG. 52 shows the velocitydistributions within the device after 500 seconds. The units are inmeters/second (m/s). It is noted that here, the inlet is from theleft-hand side of the figure, and the outlet is on the right-hand sideof the figure (flow direction indicated by the arrow). As expected, thehighest velocities are found in the inflow passageway into the acousticchamber. Negative velocities in the outflow passageway indicate flow outof the acoustic chamber. The velocity is near zero in the acousticchamber, and near the collection port at the top thereof. This velocityprofile is desirable for perfusion operation.

A computational fluid dynamics (CFD) model was also made of the devicewith the internal structure of FIG. 46C. FIG. 53 shows the velocitydistributions within the device at a flow rate of 250 mL/min, while FIG.54 shows the velocity distributions within the device at a flow rate of1000 mL/min. In these two figures, the inlet is from the right-hand sideof the figure, and the outlet is on the left-hand side of the figure(flow direction indicated by the arrow). Again, the velocity is nearzero in the acoustic chamber, and near the collection port at the topthereof, even at the much higher flow rate of 1000 mL/min. This velocityprofile is desirable for perfusion operation.

Two further CFD models were made of variants of the configurations seenin FIG. 53 and FIG. 54, which are potential internal structures for thedevice of FIG. 43. The first variant is seen in FIG. 55. Here, therectangle at the top indicates the location of the ultrasonictransducer/multi-dimensional acoustic standing wave. Below thisrectangle, the sides of the acoustic chamber taper evenly down to theoutflow passageway. The inflow passageway has an arcuate top into theacoustic chamber. The second variant is seen in FIG. 56. This variant ismuch taller and narrower compared to FIG. 55. The outflow side of theacoustic chamber tapers evenly down to the outflow passageway. As seenin both figures, the desirable velocity profile is present, being nearzero in the area of the multi-dimensional acoustic standing wave, andnear the collection port at the top thereof. These flow paths of FIGS.52-56 demonstrate how the fluid may be managed through differentconfigurations of the geometry leading from the main recirculation path(through the recirculation pipe) to the acoustic chamber. In FIG. 55 andFIG. 56, the main recirculation path to the fluid chamber have the sameflow rate. One consequence of these configurations is that theseparation velocities will be higher.

Next, the device with the internal structure of FIG. 46A was tested atdifferent flowrates. FIG. 57 shows the cell retention (%) versus time.Higher values are more desirable, and most values were over 97% atflowrates ranging from 0.37 mL/min to 1.5 mL/min. FIG. 58 shows theperfusate cell density (million cells/mL) versus time. Here, lowervalues are more desirable (indicating successful cell separation). Asexpected, results were better at lower flowrates.

The device was then tested using two different operating frequencies forthe ultrasonic transducer, 1 MHz or 2 MHz, and at different flowrates.FIG. 59 shows the cell retention (%) versus time. Higher values are moredesirable. At 2 MHz, the values were close to 100% for flow rates of 1.5mL/min to 3 mL/min. At 1 MHz, the values stayed over 90% for flow ratesof 1.5 mL/min to 3 mL/min. The frequency of 2 MHz generally performedbetter. FIG. 60 shows the perfusate cell density (million cells/mL)versus time. Again, results were better for 2 MHz operating frequency.

Example 7

Next, a device with an internal structure similar to FIG. 10 was used toperform acoustic separation of beads and yeast from a host fluid. Thedevice was a 1 inch by 1 inch by 1 inch device using a 1 inch by 1 inchcrystal to separate beads from a yeast mixture having a concentration of1×10⁶ cells/mL. The ultrasonic transducer was operated at 1 MHz with aninput voltage of 60 volts, and the beads were Solohill microcarriers inthe amount of 7.5 g/L. The mixture was flowed into the device at a flowrate of 75 mL/minute, while the concentrate was recovered at a rate of 5mL/minute. FIG. 62 shows a photograph of the microcarriers being trappedin the acoustic standing wave within the device.

Example 8

Next, a device with an internal structure similar to FIG. 7 was used toexamine AWS cell separation efficiency. FIG. 64 shows the viable celldensity (left-side y-axis), AWS power (right-side y-axis), and harvestto recirculation ratio (right-side y-axis) over a period of about 28days. The dark dots represent the bioreactor, while the light dotsrepresent the AWS harvest. The system was operated with AWS beginning onday 7 through the end of the study. From day 7 to day 20, the system wasoperated at 30% AWS power, before being switched to 10% AWS power fromday 20 to the end of the study. The harvest to recirculation ratio wasabout 15% from day 7 (start of AWS) to day 11, about 30% from day 11 tothe end of the study.

The cell separation efficiency was dominated by the harvest torecirculation flow ratio, as shown in FIG. 65 which compares the AWSpower (x-axis) to the harvest to recirculation flow ratio (y-axis). Thedevice was then paired with a 10 L bioreactor, and the AWS cellseparation efficiency was again examined. FIGS. 66A-66C show similarvolumetric productivity, improved product transmission, and improvedproduct residence time for AWS versus TFF. FIG. 67 shows the cellseparation efficiency (right-side y-axis), the harvest VCD (left-sidey-axis), and bioreactor VCD (left-side y-axis) over a period of about 14days. The system had a greater than 90% cell separation at 30% AWS powerand a 1.25% harvest to recirculation ratio.

Overall, AWS separation was shown to not affect the quality between thebioreactor and the harvest, as shown in FIGS. 68A-68C. FIG. 68Agraphically shows the relative abundance after size exclusionchromatography (SEC) of high molecular weight (HMW) species, monomerspecies, and low molecular weight (LMW) species of a monoclonal antibody(mAb). The darker bars represent the bioreactor, and the lighter barsrepresent the harvest. FIG. 68B graphically shows the relative abundanceafter ion exchange chromatography (IEX) of acidic, main peak, and basicspecies of a mAb. The darker bars represent the bioreactor, and thelighter bars represent the harvest. FIG. 68C shows glycosylation of theprotein. The effects are the same for the protein that is from thebioreactor and for the protein that has passed through the acousticseparation device. These results show that acoustics do not affect theprotein.

Finally, AWS separation was shown to have comparable quality to a TFFperfusion process using a 0.22 micron membrane, as shown in FIG. 69A andFIG. 69B. FIG. 69A graphically shows the relative abundance after sizeexclusion chromatography (SEC) of HMW species, monomer species, and LMWspecies of a monoclonal antibody (mAb). The darker bars represent AWS,and the lighter bars tangential flow filtration (TFF). FIG. 69Bgraphically shows the relative abundance after ion exchangechromatography (IEX) of acidic, main peak, and basic species of a mAb.The darker bars represent AWS, and the lighter bars represent TFF. Ascan be seen in FIG. 69A and FIG. 69B, there was a negligibly higher HWMspecies in AWS and similar charge variance profile between AWS and TFF.

Example 9

FIG. 70 depicts another device 7000 that can be used foracoustophoresis.

Device 7000 includes a first inlet port 7002 along a front wall 7011 ofthe device near the top end 7001 of the device. A second inlet port (notshown) can be located along a rear wall 7013 of the device opposite thefront wall 7011 thereof. A viewing window 7050 made of a transparentmaterial is also present in the front wall 7011. Through the viewingwindow 7050, it can be seen that an ultrasonic transducer 7020 is alsocoupled to the rear wall 7013. In this embodiment, the viewing window7050 acts as a reflector to generate the multi-dimensional acousticstanding wave(s). Both the reflector and the ultrasonic transducer aregenerally located below the inlet(s), such that the acoustic chamber ispositioned below the inlet(s). In particular embodiments, the first andsecond inlets are located directly opposite one another in the device. Adrain outlet 7006 is located near a bottom end 7003 of the device abovethe acoustic chamber, and here is located along a first sidewall 7015 ofthe device. The drain outlet generally serves as the outlet throughwhich the concentrate fluid (retained material prevented from passingthrough the standing wave) is recovered from the device. A permeateoutlet 7007, which generally serves as the outlet through which thepermeate (material that passes through the standing wave) is recoveredfrom the device, is located above the acoustic chamber, and asillustrated here is also located along the first sidewall 7015 of thedevice above the drain outlet 7006.

The device 7000 includes a cooling unit 7060 coupled to the rear wall7013. The cooling unit 7060 includes an independent flow path that isseparate from the flow path through the device containing the fluid thatis to be exposed to the multi-dimensional acoustic standing wave. Acoolant inlet 7062 is adapted to permit the ingress of a cooling fluidinto the cooling unit. A coolant outlet 7064 serves as the outletthrough which the coolant and waste heat exit the cooling unit. Here,the coolant inlet is located below the coolant outlet, though this pathcan be varied as desired. The coolant that flows through the coolingunit can be any appropriate fluid. For example, the coolant can bewater, air, alcohol, ethanol, ammonia, or some combination thereof. Thecoolant can, in certain embodiments, be a liquid, gas, or gel. Thecoolant can be an electrically non-conductive fluid to prevent electricshort-circuits. The cooling unit can be used to cool the ultrasonictransducer, which can be particularly advantageously when the device isto be run continuously with repeated processing and recirculation for anextended period of time (e.g., perfusion). The cooling unit can also beused to cool the host fluid running through the device 7000, if desired.

The performance of the device 7000 of FIG. 70 was tested using Jurkat Tcells and compared to a device having a structure similar to device 900of FIG. 9. The performance of each device is summarized in the tablebelow.

Feed (10⁶ Trapped Collected Volume Cell Device cells/mL) (%) (%)Reduction Concentrating FIG. 9 0.90 65 49 5.7 2.8 FIG. 9 0.76 84 74 2.61.9 FIG. 70 0.77 86 71 11.9 8.4 FIG. 70 0.70 75 72 11.3 8.2 FIG. 70 0.7787 77 11.4 8.9

The performance of the device was then tested using CHO cells and againcompared to a device having a structure similar to device 900 of FIG. 9.The performance of each device is summarized in the table below.

Feed (10⁶ Trapped Collected Volume Cell Device cells/mL) (%) (%)Reduction Concentrating FIG. 70 1.11 91 72 12.8 9.2 FIG. 70 1.07 90 7012.5 8.8 FIG. 70 0.99 92 67 13.0 8.7

The present disclosure has been described with reference to exemplaryembodiments. Modifications and alterations may occur to others uponreading and understanding the preceding detailed description. It isintended that the present disclosure be construed as including all suchmodifications and alterations insofar as they come within the scope ofthe appended claims or the equivalents thereof.

1. A method for separating a desired target material from a host fluid,the method comprising: flowing a mixture containing the host fluid,beads, and the target material through an acoustophoretic device, theacoustophoretic device comprising: a flow chamber including at least oneinlet and at least one outlet; at least one ultrasonic transducercoupled to the flow chamber, the at least one ultrasonic transducerincluding a piezoelectric material; and a reflector located oppositefrom the at least one ultrasonic transducer; and driving the at leastone ultrasonic transducer to create a multi-dimensional acousticstanding wave in the flow chamber; wherein the beads attach to thetarget material and are held back by the multi-dimensional acousticstanding wave; and wherein the host fluid passes through themulti-dimensional acoustic standing wave.
 2. The method of claim 1,wherein the at least one outlet includes (i) a permeate outlet throughwhich the host fluid exits the flow chamber; and (ii) a concentrateoutlet through which the target material exits the flow chamber.
 3. Themethod of claim 1, wherein the beads are not functionalized.
 4. Themethod of claim 1, wherein the beads are functionalized.
 5. The methodof claim 1, wherein the beads have a positive contrast factor.
 6. Themethod of claim 5, wherein the beads are selected from the groupconsisting of polystyrene beads and glass beads.
 7. The method of claim1, wherein the beads have a negative contrast factor.
 8. The method ofclaim 7, wherein the beads are selected from the group consisting ofmicrobubbles and micro-glass spheres.
 9. The method of claim 1, whereinthe beads are polymeric.
 10. The method of claim 1, wherein the beadsare glass, hollow, or gas-filled.
 11. The method of claim 1, wherein thebeads are spherical, toroidal, cylindrical, or conical.
 12. The methodof claim 1, wherein the desired target material is microvesicles,viruses, proteins, recombinant proteins, or monoclonal antibodies. 13.The method of claim 12, wherein the microvesicles are exosomes oroncosomes.
 14. The method of claim 1, wherein a pressure rise and anacoustic radiation force on cells are generated at an interface regionto clarify the host fluid as the mixture passes through themulti-dimensional acoustic standing wave.
 15. The method of claim 1,wherein the acoustophoretic device includes a cooling unit for coolingthe at least one ultrasonic transducer.
 16. The method of claim 1,wherein at least 95% of the cells pass through the multi-dimensionalacoustic standing wave.
 17. An acoustophoretic device, comprising: aflow chamber including at least one inlet port and at least one outlet;at least one ultrasonic transducer coupled to the flow chamber and atleast one reflector opposite the at least one ultrasonic transducer; anda cooling unit for cooling the at least one ultrasonic transducer.